Particle suitable for the manufacture of an implantable soft tissue engineering material

ABSTRACT

The particle ( 1 ) is suitable for the manufacture of an implantable soft tissue engineering material and comprises a three-dimensionally warped and branched sheet ( 2 ) where: 
     (i) the three-dimensionally warped and branched sheet ( 2 ) is made from a biocompatible material having a Young&#39;s modulus of 1 kPa to 1 GPa; 
     (ii) the three-dimensionally warped and branched sheet ( 2 ) has an irregular shape which is encompassed in a virtual three-dimensional envelope ( 3 ) having a volume VE; 
     (iii) the three-dimensionally warped and branched sheet ( 2 ) has a mean sheet thickness T; 
     (iv) the three-dimensionally warped and branched sheet ( 2 ) has a volume VS; 
     (v) the particle ( 1 ) has a Young&#39;s modulus of 100 Pa to 15 kPa; and 
     (vi) the particle ( 1 ) further comprises a number of protrusions ( 4 ) where the three-dimensionally warped and branched sheet ( 2 ) reaches the envelope ( 3 ); 
     (vii) the particle ( 1 ) has a number of interconnected channel-type conduits ( 5 ) defined by the branching of the sheet ( 2 ) and/or by voids in the sheet ( 2 ); and 
     (viii) where the conduits ( 5 ) have (a) a mean diameter DC; and (b) an anisotropicity index of 1.01 to 5.00.

BACKGROUND OF THE INVENTION 1. Field of the Invention

The invention relates to a particle suitable for the manufacture of animplantable soft tissue engineering material, to an implantable softtissue engineering material comprising a multitude of particles andmethod for manufacturing such particles.

2. DESCRIPTION OF THE RELATED ART

From U.S. Pat. No. 7,993,679B2 a flowable collagen/glycosaminoglycanmaterial is known which can be used as a wound dressing. This materialhas no capacity to be expandable.

From U.S. Pat. No. 5,629,191 a method is known for making a mass ofgeneral spherical porous matrix particles, which, however, have nocapacity to undergo expansion upon uptake of liquid and have sphericalshapes with no protrusions.

From the scientific article “Injectable preformed scaffold withshape-memory properties”, Bencherif et al., Proceedings of the nationalacademy of sciences, vol. 109, no. 48, Nov. 12, 2012, pages 19590-19595a process is known in which a polymerizable material is cross-linkedusing a radical polymerization process starting at a temperature of +4°C. and continued at an unknown temperature. The premixed polymerizationmixture is then poured on a pre-cooled mold producing a skinning effectconsisting of the partial closing of external pores. The polymerizedscaffold material is then used as such for injection into a patient,i.e. with no further fractioning of the scaffold obtained.

Other prior art manufacturing processes start from pre-existing tissuesas a raw material for producing particulate implantable soft tissuematerials. The disadvantage of such methods lays in the fact thathardness and elasticity of the particles, which are essential for theintended volume reconstruction process, cannot be adjusted.

Other known processes start from frozen material, which gives lesselastic and less porous particles.

A) General Definitions

Soft Tissue Engineering Material

A biocompatible material for use in reconstruction, repair, enhancement,augmentation or substitution of soft tissues. In the context of thisinvention, a material obtained by combination of the branched-sheet 3Dparticles with a physiologically acceptable fluid (physiological saline,phosphate buffered saline, blood plasma, living aspirates such aslipoaspirate).

Dry Mass Concentration

The dry mass concentration of the soft tissue engineering material canbe determined by standard techniques known in the art, i.e. washing ofthe particles to remove salts followed by drying to constant weight, andrelating the dry weight found in this way to the volume of the softtissue engineering material.

B) Definitions Concerning the Reference Conditions

Reference liquid and reference measurement condition The geometric andmechanical properties of sheet material, particles, and soft tissueengineering material all depend on the suspension liquid. It istherefore necessary to specify the reference liquid in which such valuesare measured. For the present invention, this liquid is phosphatebuffered saline, with the following concentrations: potassium chloride(KCl) 2.6 mM, potassium phosphate monobasic (KH₂PO₄): 1.5 mM, sodiumchloride (NaCl) 138 mM, sodium phosphate dibasic (Na₂HPO₄): 8.1 mM indeionized water. Knowing these concentrations, a person skilled in theart will be able to prepare the reference liquid. The measurement ofgeometrical and mechanical properties of samples should be carried outwith samples immerged for at least 24 h with at least a 100× excess ofreference liquid relative to the sample volume prior to the measurementand immerged in the reference liquid during the measurement. Thereference temperature is room temperature (25° C.).

C) Geometric Definitions

Particle Size

The size of a particle is defined by the volume of its convex hull. Theconvex hull is a well-known mathematical concept; it comprises all thematerial points making up the particle, and also all possiblestraight-line connections between these material points. The convex hullcan be determined by confocal imaging of freely suspended, isolatedparticles in the reference liquid. A non-exclusive example ofdetermination of the size of a particle in accordance with the inventioninvolves the following steps, known to a person skilled in the art:

1) Confocal microscopy according to techniques known in the art is usedto acquire an image of an isolated particle freely suspended in thereference liquid. Confocal imaging can typically be done based onautofluorescence of the sheet material or fluorescence from dyesdisplaying affinity or on the contrary being excluded from the sheetmaterial. 2) The confocal images acquired are stored electronically asstacks of images, each image corresponding to a confocal plane. Thepixels making up these images are also referred to as voxels since theyrepresent the fluorescence information from a confocal volume of acertain depth.

3) Each voxel is then either assigned to the sheet material or the voidspace (based on thresholding between two very distinct populations ofintensities).

4) Using this binarized representation, the convex hull constructedusing available software. For the particles at hand, the convex hulldefines a mostly polyhedral 3D volume that contains all the voxels ofthe branched sheet but also the pore space interior to the particle.

5) Finally, the total particle volume can be calculated by the number ofvoxels identified as belonging to the convex hull times the volumeassociated with each voxel.

Equivalent Particle Diameter

The equivalent particle diameter is the diameter of sphere that has thesame volume as the particle's convex (the “particle size”). To bespecific, if the volume of the particle is V, then the equivalentdiameter is given d=(6*V/pi)∧(⅓), where pi=3.1419 . . .

Envelope of a Particle

The envelope of a particle is defined by the surface of its convex hull.As a non-exclusive example, a digital representation of the convex hullof a given particle can be obtained by the sequence outlined for theparticle size above (i.e. confocal images of freely suspended particlesin the reference fluid are acquired, stored, and analyzed as describedfor the size of the particle to define its convex hull). The envelope ofthe particle is defined by the voxels of the convex hull that have atleast one direct neighbor (sharing a face or corner) that does notbelong to the convex hull.

Protrusions of the Particle

The protrusions with the particle surface are defined as the parts ofthe sheet that reach the envelope of the particle as defined above. As anon-exclusive example, a digital representation of the protrusions canbe defined from the approach described for the envelope of the particle.Indeed, once the digital representation of the envelope is defined, thelocation of the waU protrusions is found by taking the intersection ofthe voxels representing the branched sheet and the voxels representingthe particle envelope.

Protrusion Depth

The protrusion depth is defined by the length of a normal vector from apoint of the particle envelope to the closest intersection point with asheet of the particle. If the point of interest of the particle envelopbelongs itself to a sheet, the associated protrusion depth is zero. Ifthe normal vector has no intersection with a sheet, the distance betweenthe two intersection points of the normal vector with the envelope istaken instead.

Maximal Protrusion Depth

The maximal protrusion depth is largest observed protrusion depth for agiven particle.

Relative Maximum Protrusion Depth

The relative maximum protrusion depth is defined as the maximumprotrusion depth divided by the equivalent particle diameter. If largerthan 1, a value of 1 is recorded.

Contact Area of the Protrusions

The contact area of the protrusions is the area where the distance fromthe envelope to the nearest sheet is less than the average sheetthickness TC.

Conduits

The conduits are the voids inside the particles. In the soft tissueengineering material, they can be filled with a physiologicallyacceptable fluid.

Conduit Diameter D_(C)

The conduit diameter D_(C) is defined as the mean Feret diameter of theconduits as individual void areas on a section of the particle. Todetermine the conduit diameter D_(C), 2D cross section images ofparticles freely suspended in reference liquid may be obtained, forinstance by confocal imaging. On such sections, the conduits areidentified as the visible distinct voids; for conduits reaching theparticle exterior on the section of interest, the boundary of the convexhull of the particle as visible on the 2D section is used to completedelimitation of the void by the sheet material. For each of the voidsidentified, the Feret diameter, known to a person skilled in the art, ismeasured in all possible orientations (in practice, a few hundred), andthe average taken for each void. To obtain the average conduit diameter,the average of these averages is taken, ideally for several sections ofseveral particles.

Thickness of the Sheet T

The average thickness of the sheet T is defined as the mean thickness ofsheet material encountered when randomly slicing the particle. Todetermine the average thickness of the sheet T, 2D cross section imagesof particles freely suspended in reference liquid may be obtained, forinstance by confocal imaging. On such images, pixels are eitherattributed to the sheet material or the pore fraction, typically basedon intensity thresholding. To determine the sheet thickness, randomstraight lines are drawn across the particle (random orientation, randomlocalization). For each such line, transitions from void to sheet andback are noted, each such transition from void to sheet and backrepresenting a sheet intersection. The local thickness of the sheet forsuch a sheet intersection is determined from the number of pixelsbelonging to the walls and the known size of the pixels. The averagethickness of the sheet T is obtained by averaging over all the sheetintersection, ideally from hundreds of lines per particle and also astatistically relevant number of particles.

Relative Sheet Thickness

The relative sheet thickness is the ratio of the sheet thickness toconduit diameter: T_(relative)=T/(D_(C)+T). For particles with lowrelative sheet thickness T_(relative) (i.e. T_(relative)<0.1) oneexpects the total sheet volume fraction to be proportional toT_(relative), with a proportionality constant near unity.

Conduit Diameter, Sheet Thickness and Relative Sheet Thickness in theSoft Tissue Engineering Material

Conduit diameter, sheet thickness and relative sheet thickness are notonly defined for the isolated particles, but also for the soft tissueengineering material. In this case, they are understood to depend on theamount and nature of the fluid added to obtain the soft tissueengineering material. They can be measured by the same methodology asthe one exposed for the particles, i.e. by acquisition of planar imagesfor instance by confocal imaging, following by the image analysistechniques defined above. In the case of the soft tissue engineeringmaterial, one should be aware that the conduits detected as voids willgenerally be conduits within particles, but also conduits formed betweenneighboring particles; except for if particular precautions such aslabeling of different particles with different dyes are taken, it is notgenerally possible to distinguish between the two. Likewise, some of thesheet intersections in the procedure for determining sheet thicknesswill be formed by touching sheets from neighboring particles, withoutthis being necessarily evident in image analysis.

Branched Sheet Morphology

The branched sheet morphology applies to the particles. Porous materialsin general are classified either as “open” or “closed” foams (CellularSolids: Structure and Properties, Gibson and Ashby). In the terminologyof foam science, closed foams have “cells” (pores, voids) where thefaces shared with neighboring “cells” are solid membranes (CellularSolids: Structure and Properties, Gibson and Ashby); these closed foamsentrap the pore fluid, which cannot easily escape through the solidmembranes delimiting the “cells”. Classical open foams have no membranesbetween neighboring cells, but only struts along the edges where threeor more cells meet (Cellular Solids: Structure and Properties, Gibsonand Ashby). Such open foams naturally have a fully interconnected porespace, and a fibrous network of solid material. The branched sheetmorphology can be seen as a special structure that functionally combinesaspects of open and closed foams. In a perfect branched sheetmorphology, each “cell” can be understood to have both open and closedfaces. The open faces provide connections to neighboring “cells”, whilethe closed faces make up the bulk of the branched sheet. The open facesare organized in such a way to allow connectivity to the particleexterior for the vast majority of the foam cells, and hence to providepore interconnectivity. The closed faces together provide the branchedand warped 3D-sheet morphology. Typically, each “cell” has at least twoopen and two closed faces, where an open face can also be a connectionto the particle exterior.

Anisotropicity Index

The anisotropy index is the ratio between the longest Feret diameter tothe smallest Feret diameter of a body or void. In the context of thisinvention the anisotropicity index refers to the conduits, i.e. theratio between longest and smallest Feret diameter for each conduit, asindividualized as a distinct void in 2D cross-sectional images (seeconduit diameter D_(C)). The global anisotropy index is the average ofthe anisotropy indexes associated with the individual conduits, asidentified as individual voids on 2D cross-sections.

D) Definitions Concerning Mechanical Properties

Younq's Modulus

The Young's modulus of a given material is the ratio between stress(force per area) and strain (deformation relative to the originallength) as measured in uniaxial compression, typically for small strains(10% compression or less). For the present invention, Young's moduli aredefined at three different scales: the Young's modulus of the materialfrom which the branched sheets are made; the Young's modulus of theresulting particles; and the Young's modulus of the assembled softtissue engineering material.

Young's Modulus of the Sheet Material

The Young's modulus of the sheet material is the Young's modulus asmeasured on a pure, non-porous sample of the sheet material, in thereference liquid. In some embodiments, the sheet material can not onlybe synthesized as porous, branched sheet particles, but also assufficiently large bulk material. In this case, the most precise methodof evaluation of the Young's modulus of the sheet material is by directmechanical testing of a bulk sample of the sheet material (underreference conditions). In other embodiments, it is difficult to producethe sheet material in homogeneous bulk quantities; in this case, theYoung's modulus of the sheet material can still be evaluated bymicroindentation techniques, directly on structured materials, as forinstance described in Welzel et al., Adv Healthc Mater 2014, or Beduerat al., Advanced Healthcare materials 2015. 4(2): p. 301-12, using thereference conditions given above.

Young's Modulus of the Particles

As there is no technique known to directly assess the Young's modulus ofthe irregular, porous particles of the present invention, macroscopicsamples with regular geometries (typically cylinders of a few mm heightsand 1 or 2 cm diameter), but with identical pore structure and sheetmaterial are produced for mechanical testing. It is then possible to usestandard mechanical testing equipment to impose the desired uniaxialdeformation, and to measure the associated force to determine strain(deformation relative to original sample height) and stress (force percross-sectional area) to calculate the Young's modulus of the particlesunder reference conditions. As it will be known to a person skilled inthe art, the mechanical testing of porous samples heavily depends on theboundary conditions for the pore fluid. Under the reference conditions,the sample is immersed in reference fluid, and is therefore underso-called drained conditions. Nevertheless, to avoid undue influence ofthe drainage of the pore fluid on the measured Young's modulus, it isimportant to ensure that the fluidic drainage resistance of the samplebecomes negligible with respect to the solid's response. In practice, itwill be known to a person skilled in the art that the primary variablesto control in this respect are sample geometry and compression rate, andthat suitable compression conditions can be found for instance throughminimization of hysteresis or validation by stress relaxation tests.

Young's Modulus of the Soft Tissue Engineering Material

A third Young's modulus applies to the soft tissue engineering material.The soft tissue engineering material consists of a given amount ofparticles (specified for instance relative to the original amountfabricated or absolutely in terms of dry mass present), and a specifiedamount of reference fluid. For a wide range of fluid volume to particlemass, this mix forms a paste with a non-zero yield stress and strain(measurable through rheometry), meaning that the paste behaves like asolid for sufficiently small deformations. For such small deformations,the Young's modulus can be measured by uniaxial compression. The Young'smodulus naturally depends on the ratio of particles to fluid, thecomposition and geometry of the particles, as well as the composition ofthe fluid.

In practice, it should be noted that soft tissue engineering materialcannot simply be immerged into the reference fluid, undue dispersal ofthe particles would result. Instead, the drainage of reference porefluid during compression has to be carried out through a porousmembrane, with a pore size such that the reference fluid easily passes,but the particles are retained in the compression area. A person skilledin the art will easily be able to adapt existing setups to this purposeor identify a suitable commercial apparatus.

Deployment Pressure

The deployment pressure is the pressure with which the soft tissueengineering material of this invention expands, and therefore attractsliquid. Like the Young's modulus of the soft tissue engineeringmaterial, the deployment pressure depends on the exact nature of theparticles suspended, the exact composition of the fluid used for itsmeasurement, and the ratio of particles to fluid. The deploymentpressure of the soft tissue engineering material can be measured bycontacting the soft tissue engineering material with a mesh or membranefreely permeable to the reference fluid, but not the particles of thesoft tissue engineering material. The deployment pressure is then thedifference between the pressure recorded in the soft tissue engineeringmaterial and the pressure in pure reference fluid in equilibrium withthe soft tissue engineering material. A possible readout technique isobservation of hydrostatic height differences in a U-tube configurationwith the separation membrane between soft tissue engineering materialand the pure reference fluid; another possibility is the use of pressuregauges, one where only the reference fluid has access, the other exposedto the soft tissue engineering material.

Maximal Elongation of the Sheet Material

The maximal elongation of the sheet material is the maximum relativedeformation of the sheet material compatible with essentially purelyelastic deformation in the reference liquid: beyond this maximalelongation, the sheet material fails by either onset of plasticdeformation or brittle fracture.

Swelling Value of a Material

The swelling value is defined as the relative weight change of amaterial from dry to humid state. This parameter is a ration, namely(Humid weight−Dry weight)/Dry weight. The measurement of the swellingvalue is usually performed in physiological saline 0.9% NaCl.

BRIEF SUMMARY OF THE INVENTION

It is an object of the invention to provide a particle suitable for themanufacture of an of implantable soft tissue engineering material withthe capability of the particles to increase the size of its conduits byfluid uptake and therefore to increase the volume of the particles up toa predefinable and controllable volume.

This is achieved thanks to the particular properties of the branchedsheets composing the particle.

The invention solves the posed problem with a particle as disclosed andclaimed herein and with an implantable soft tissue engineering materialas disclosed and claimed herein.

The advantage(s) of the particles according to the invention and of theimplantable soft tissue engineering material comprising such particlesis to be seen in the ability of the particles to absorb water or aqueoussolutions, like body fluids, thereby causing the three-dimensionallywarped and branched sheet to deploy so that its envelope will occupy alarger volume. This volume is predefined by the conduits volume.

Compared to other materials the particles according to the inventionhave shown the unique property that they are capable of deploying backgently, yet rapidly under in-vivo conditions. The reason for this is tobe seen in the deployment pressure required for partial swelling thatmatches the in vivo physiological deployment pressure. For example, toosoft porous materials flow and cannot sustain a three-dimensionalvolume. Since these known materials are too soft its pores (or conduits)are prevented to actively deploy. On the other hand, too hard materialsprovide a foreign body not matching the mechanical properties of thesoft target tissues and are creating a local mismatch of Young'smodulus, at the origin of a local inflammatory reaction and possiblyfibrosis.

The capacity of deployment of the branched sheet of the particleaccording to the invention combined with the interconnected porosity inthe form of channel-like conduits enables a better bio-integration, i.e.ingrowth of tissues and vessels into the particle compared to state ofthe art particle without the capability of deployment.

The injectable material allows the filling of tissue defects of smalland medium size and even defects larger than 50 cm³.

The protrusions of the particles—which produce an unexpectedzip-fastener effect—are an essential feature of the invention. Thisfeature enables cohesivity of the engineering material implanted intothe patient. The protrusions are a key feature enabling cohesion betweenthe particles and avoid migration/permanent deformation of theengineering material after its injection into the patient.

A further essential feature of the invention is the degree ofanisotropicity of the conduits enabling the volume of the engineeringmaterial to be maintained.

It has been found that if the conduits are too anisotropic, the sheetswill collapse and stick to each other, leading to spreading of theimplanted engineering material and preventing creation of a stablethree-dimensional volume.

The deployment of the branched sheets of the particles enables a betterand rapid ingrowth of cells, tissues, blood vessels, lymph andextracellular matrices and more stability of the shape and volume of theimplanted engineering material.

The zipper-effect of the protrusions enables to maintain the injectedshape by avoiding migration of the particles.

Thanks to the combination of the above mentioned features and to theactive deployment, the engineering material is able to enter into africtional relationship with the surrounding tissues enabling theengineering material to stay in place (anchoring).

If the implanted material is too hard, this anchoring effect is replacedby a cutting effect and the tissue is damaged. If the material is toosoft, the material is not anchoring and slides over the tissues.

Deployment at low pressures enables the gentle aspiration of tissues orcell suspensions for co-grafting applications, for example mixing of theengineering material with adipose tissues and injection of the mixtureto create a living volume.

The low deployment pressure results from particle design. Theapplication imposes the Young's modulus of the particles, but dependingon the particle structure, the deployment pressure can vary. To obtain alow deployment pressure for a given Young's modulus, it is essentialthat compression of the particles can occur as much as possible byevacuation of pore fluid, and not by compression of the incompressiblepore fluid or nearly incompressible sheet material itself, as thisnecessitates greater isotropic pressure. For this, the branched sheetmorphology is key: the pore fluid can easily be evacuated throughinterconnectivity, whereas the closed faces provide mechanical solidity,allowing to keep the sheet volume fraction low and avoidingsheet-to-sheet contact and compression. This combination gives thedesired low deployment pressure for the desired Young's moduli.

In a special embodiment of the invention the volume V_(E) of theenvelope is larger than 5-10⁻⁴ mm³, preferably larger than 4-10⁻³ mm³and most preferably larger than 0.03 mm³ The volume V_(E) of theenvelope is purposefully smaller than 4-10³ mm³, preferably smaller than260 mm³, and most preferably smaller than 30 mm³.

The thickness T of the sheet 2 is purposefully larger than 1 μm,preferably larger than 5 μm, The thickness T is purposefully smallerthan 1000 μm, preferably smaller than 100 μm, and most preferablysmaller than 20 μm. Surprisingly it has been found that a sheetthickness in that range allows obtaining a material which is transparentwhen imaged with ultrasounds imaging devices (such as the ones used inpatients.

The total surface of the sheet 2 is purposefully larger than 8-10⁻⁴ mm²,preferably larger than 1-10⁻² mm², and most preferably larger than 10⁻¹mm². The total surface of the sheet 2 is purposefully smaller than 50000mm², preferably smaller than 10000 mm².

The total volume V_(S) of the sheet 2 is purposefully larger than 8-10⁻⁷mm³, preferably larger than 5.10-10⁻⁶ mm³, and most preferably largerthan 5-10⁻⁴ mm³. The total volume V_(S) of the sheet 2 is purposefullysmaller than 50000 mm³, preferably smaller than 5000 mm³, and mostpreferably smaller than 200 mm³.

In a special embodiment the anisotropicity index of the conduits 5 islarger than 1.05, preferably larger 1.10. The anisotropicity index ispurposefully smaller than 10.0, preferably smaller than 5.0 and mostpreferably smaller than 3.0.

The material for the sheet can be chosen from non-fouling substancespoly-ethyleneglycol (PEG), poly-acrylamide,poly-(Hydroxyethyl)methacrylate, preferably from polysaccharides such ascellulose, methylcellulose, carboxymethylcellulose, agarose, polysucroseor dextran.

The sheet may also comprise a material chosen from the following group:carbohydrates, hydrogels, collagens, gelatins, peptides or extracellularmatrices. These materials confer biocompatibility and biodegradability.

The Carbohydrate can be

(i) a polysaccharide, preferably a negatively charged polysaccharide;

(ii) an alginate;

(iii) hyaluronic acid; or

(iv) a carboxymethyicelluose.

These materials confer elasticity to the particle enabling a reversiblecompression capability and a reversible fluid intake capability Thesheet may also comprise a synthetic polymer, preferably chosen from thefollowing groups

(i) silicones;

(ii) polyurethanes;

(iii) polyolefins;

(iv) acrylates, preferably poly-acrylamide or poly-acrylic acid;poly-(Hydroxyethyl)-methacrylate and copolymers thereof;

(v) polyesters;

(vi) polyamides; or

(vii) polyimides.

The material of the sheet 2 has preferably a maximal elongation of VS/VEor more. This confers a reversible fluid intake capability to theparticle.

The material of the sheet 2 has preferably a maximal elongation of3T/D_(C) or more.

The material of the sheet 2 has purposefully a molecular weight in therange of 50 Da-10 MDa.

In a further embodiment the ratio of D^(C)/T is larger than 1.0,preferably larger than 2. Purposefully the ratio of D^(C)/T is largerthan 5, preferably larger than 7. The ratio of D^(C)/T may be smallerthan 500, preferably smaller than 100. Purposefully the ratio of D^(C)/Tis smaller than 50, preferably smaller than 30. The ratio of D^(C)/T maybe smaller than 29, preferably smaller than 15. The specific choice ofthe ratio of D^(C)/T combined with the choice of material enables thereversible compressibility of the particle and obtaining suitableparticle Young moduli.

In a further embodiment the mean diameter D_(C) of the conduits islarger than 1 micrometer, preferably larger than 14 micrometer.Purposefully the mean diameter D_(C) of the conduits is larger than 20micrometers, preferably larger than 50 micrometers. The minimumdimensions for the conduits enable cellular and vascular ingrowth intothe particle.

In a further embodiment the mean diameter D_(C) of the conduits issmaller than 10 mm, preferably smaller than 4 mm. The mean diameterD_(C) of the conduits may by smaller than 2 mm, preferably smaller than1 mm.

Purposefully the mean diameter D_(C) of the conduits is smaller than 600micrometer preferably smaller than 300 micrometers. The maximumdimension for the conduits enables a mechanical stability of thescaffold and tight contacts between the scaffold and the tissues. If thevalue for D_(C) is too high, the cellularization and vascularization isnot as high as if D_(C) is optimal.

It was surprisingly found that the choice of these parameters allowsobtaining a material which matches the speckled appearance of nativetissue when imaged with ultrasound imaging devices (such as the onesused in patients.

In a further embodiment the protrusions 4 have a mean relative maximumprotrusion depth in the range between 0.05 and 1.0, preferentiallybetween 0.15 and 0.8.

The shape of the three-dimensionally warped and branched sheet 2 ispurposefully flexible and in particular it is preferably reversiblyexpandable upon absorption or removal of a liquid by the biocompatiblematerial. Liquids suitable for absorption are water, aqueous solutions,blood or other body fluids.

In a further embodiment the Young's modulus of the biocompatiblematerial of the sheet 2 is at least 200 Pa, preferably at least 1′000Pa. Purposefully the Young's modulus of the biocompatible material ofthe sheet 2 is smaller than 500 kPa, preferably smaller than 50 kPa. Thematerial constituting the sheets should have a Young's modulus highenough to avoid compression of the particle at pressures that can befound in the body (fluid pressure of the interstitial fluid). If thematerial is too “soft”, it has become apparent that the sheets collapseone against to each-other and they are not able to expand back to obtaina certain “memory shape” effect observed with stiffer sheets

In a further embodiment the mean diameter D_(P) of the particle islarger than 2 micrometers, preferably larger than 10 micrometers. Thisfeature enables the particle to create a cohesive implant. If particlesare too small, there is the risk of migration of the particle in lymphor blood vessels, or phagocytosis by macrophages for example.

The mean diameter D_(P) of the particle is purposefully smaller than 5mm, preferably smaller than 2 mm. This limitation of the particledimension enables at the same time structurability and shapeability of apaste created by a plurality of particles and a cohesivity of the paste.

In a further embodiment the particle comprises at least 5 conduits,preferably at least 10 conduits. A small number of conduits in theparticle leads to mechanical properties which are less optimal and theparticle will have more chances to collapse.

In a further embodiment the ratio between the mean diameter of theconduit and the mean diameter of the particle D_(C)/D_(P) is larger than1.5, preferably larger than 2.0 Surprisingly it has been found that anoptimal D_(C)/D_(P) ratio avoids reduced porosity which would reducealso the tissue ingrowth capability. The optimal ratio also leads to afavorable the biodegradation time. Purposefully the ratio between themean diameter of the conduit and the mean diameter of the particleD_(C)/D_(P) is smaller than 20, preferably smaller than 10.

In a further embodiment the contact angle between water and thebiocompatible material of the sheet 2 is in the range of 0° to 90°,preferably in the range of 0°-60°. This feature allows an optimal liquidintake in the conduits of the particle.

Preferably the Sheet 2 is Reversibly Compressible.

In a further embodiment the particle is hydrated and preferablycomprises at least 0.05 weight-% of the biocompatible material based onthe total weight of the hydrated particle. Purposefully the hydratedparticle comprises at least 0.1 weight-%, preferably at least 0.5weight-% of the biocompatible material.

Purposefully the particle is hydrated and comprises at most 15 weight-%of the biocompatible material based on the total weight of the hydratedparticle. The hydrated particle may comprise at most 5 weight-%,preferably at most 3 weight-% of the biocompatible material.

The particle according to the invention may comprise severalthree-dimensionally warped and branched sheets 2.

The invention therefore is also directed to a composition comprising:

a) a multitude of particles according to the invention; and

b) a physiologically acceptable fluid.

The amount of fluid may be such that the particles are only partiallyhydrated. This has the advantage that the particles of the compositionretain still the ability to deploy by up-take of body fluids afterinjection into the body of the patient.

The invention is further directed to an implantable soft tissueengineering material comprising a multitude of particles according tothe invention, preferably in form of a malleable paste. The multitude ofparticles may be admixed with water or an aqueous solution or blood toform a malleable paste. The multitude of particles may also be admixedto adipose tissue. This has the advantage that the mixing of the pastewith adipose tissues allows creating a new volume of adipose tissues.

The implantable soft tissue engineering material according to theinvention is purposefully reversibly compressible after injection into apatient by uptaking liquid from the surrounding tissues. The uptake ofliquid occurs up to a fixed predefined amount corresponding to the fullexpansion state of the particles, i.e. until the conduits are filledwith liquid. Once the conduits are full, the expansion cannot gofurther. In vivo, there is then an equilibrium between the particlespressure and the pressure of the in vivo interstitial pressure. At thisstage the material is not deformable anymore guaranteeing that thevolume created for the patient cannot be deformed anymore.

The implantable soft tissue engineering material according to theinvention purposefully

exhibits a non-linear compression behavior and a Young's moduluscomprised between 100 Pa and 15 kPa.

The implantable soft tissue engineering material according to theinvention may be used as a shapeable tissue or organ body implant. Theimplantable soft tissue engineering material according to the inventionmay also be used for treating tissue defects, in particular tissuedefects caused by severe trauma or cancer ablation. I may also be usedfor breast reconstruction and for lipofilling. Further the material maybe used for aesthetic restorations in the face and the body.

The invention is further directed to a method for manufacturingparticles according to the invention comprising the following steps:

a) pre-cooling a polymerizable biocompatible material in an aqueoussolution at a temperature below 10° C.;

b) cross-linking the pre-cooled mixture at a temperature below 0° C.,preferably below minus 1° C.; and

c) fractioning the cross-linked biocompatible material obtained.

The cross-linking according to step b) may purposefully be performed ata pH-value of minimum 5.0, preferably minimum 5.5.

The cross-linking according to step b) may purposefully be performed ata pH-value of maximum 8.5, preferably maximum 7.5. The cross-linkingaccording to step b) may be performed at a pH-value of maximum 6.9,preferably maximum 6.5.

The cross-linking process in step b) should not be one based on aradical polymerization since it has been found that radicalpolymerization can pose a hazard problem because of incorporation ofinitiator and possible depolymerization during sterilization in vitro.

The cross-linker used in step b) may be adipic dihydrazide.

In a special embodiment the cooling process in step b) consists of twosub-steps:

-   -   (i) a first sub-step to a temperature in the range of 0° C. to        −15° C., preferably in the range of −2° C. to −12° C., followed        by a hold time for temperature equilibrium; and    -   (ii) a second sub-step to a temperature in the range of −80° C.        to −2° C.

The polymerizable biocompatible material has preferably a molecularweight of 50 Da-10 MDa.

A BRIEF DESCRIPTION OF THE DRAWINGS

A special embodiment of the invention will be described in the followingby way of example and with reference to the accompanying drawings inwhich:

FIG. 1 illustrates a three-dimensional view of an embodiment of theparticle according to the invention and its virtual envelope;

FIG. 2 shows the same view as FIG. 1 highlighting the protrusions at theperiphery of the particle by means of broken-line circles;

FIG. 3 shows the same view as FIG. 1 in which the thickness T of thesheet at various locations is indicated by arrows.

FIG. 4 shows the same view as FIG. 1 in which the thickness diameter DCof some of the conduits is indicated by arrows.

FIG. 5 shows the same view as FIG. 1 in which the protrusion depth isindicated by the length of a normal vector Nv from a point of theparticle envelope to the closest intersection point with a sheet of theparticle.

FIG. 6 shows 3-D projections of a material having a Young's modulus tolow (A) and of a material according to the invention (B).

FIG. 7 is a graphical representation of the percentage of cells retainedin the soft tissue engineering material according to the invention whenseed in vitro with and without deployment effect.

FIG. 8 is a graphical representation of the percentage ofcellularization of the soft tissue engineering material according to theinvention.

FIG. 9 shows macroscopic pictures of the implants site over time.

FIG. 10 represents photographs regarding histology and macroscopicobservation of bio-integration of comparative examples.

FIG. 11 is a graph showing cellular invasion and vascularization ofcomparative examples.

FIG. 12 is a graph showing the vessels density (number of vessels/mm²)observed on histology pictures after the implantation of a soft tissueengineering material having a mean diameter of the conduits of 16micrometers and a soft tissue engineering material having a meandiameter of the conduits of 127 micrometers.

DETAILED DESCRIPTION OF THE INVENTION

The following examples clarify the invention further in more detail.

A) Manufacture of the Particles

Example 1

Carboxymethyl-cellulose (with a MW of 700 kDa) was dissolved indeionized water to the concentration of 2%, and crosslinking initiatedafter precooling to 4° C. by means of addition of adipic aciddihydrazide AAD (0.07%) and a small excess of the carbodiimide EDC(0.4%) and buffered to a pH-value of 5.5. by means of an excess of2-(N-morpholino)ethanesulfonic acid (MES) buffer (50 mM).

The reaction mixture was placed at −20° C. in a mold. After 1 day, thescaffolds were thawed and washed in de-ionized water (DI).

The next step consisted in fractioning the scaffold. For this, a bulkscaffold or a bulk scaffold piece was placed in a plastic bag andcompressed and sheared manually to create the particles according to theinvention. In another embodiment, the bulk scaffold was extruded througha thin tubular element by applying a known pressure to obtain afragmented material.

The particle size was controlled by the pressure applied on the pistonof the syringe and by the size of the extruding cannula. Typically, apressure of 15 bars and a cannula of 14G was used.

Example 2a

The same procedure as in example 1 was followed but prior to freezing,the reaction mixture was distributed into a silicone mold using apipette of 10 mL. The silicone mold contained microstructuredstar-shaped cavities measuring 100 micrometers in diameter and 20micrometers in depth. The silicone mold was covered with a flatpolypropylene counterpart, squeezing excess liquid from the mold. Theassembly was then placed into a freezer at −20° C.

Alternative Methods for the Manufacture of Particles:

Particles were manufactured by placing the scaffolds into a mixer andmixing them.

Particles were manufactured by ink-jet printing, 3D printing, andadditive manufacturing.

Particles were manufactured by mixing the reaction mixture with aphotosensitizer (typically acrylamide monomer andN,N′methylenebis(acrylamide), freezing at −20° C. and photopolymerizingusing a UV lamp or a visible lamp.

Particles were manufactured by grinding a preliminary manufacturedscaffold during at least 30 s, for example using a mixing robot (forexample Kenwood Major Titanium KMM060).

Particles were manufactured by cutting and/or slicing a preliminarymanufactured scaffold using cutting blades, possibly organized innetworks.

It is important to note that a classical emulsion polymerization methodwould give nearly perfectly round particles and therefore would not leadto the desired structure of the particles according to the inventionwith significant protrusions.

Example 2b

A solution of 5% of hyaluronic acid monomers with a molecular weight of90 kDa, MES buffer pH6, adipic acid dihydrazide (2 mg/mL) was mixed withEDC (4 mg/mL) and poured onto a consolidated paraffin microspheresscaffold. The paraffin beads were prepared according to “Microspheresleaching for scaffold porosity control”, Draghi et al, Journal ofMaterial sciences: Materials in medicine, 16 (2005) 1993-1997. Themixture was incubated at room temperature during 24 hours after whichthe paraffin beads were dissolved by an excess of hexane. The obtainedscaffold was then rinsed with isopropanol, and a mix of isopropanol andwater (40%:60%) and followed by a rinsing step with water.

The obtained scaffold was then fragmented by applying an extrusion forceon the scaffold through a narrow tubular element.

B) Manufacture of an Implantable Soft Tissue Engineering MaterialComprising a Multitude of Particles According to the Invention

Example 3

Carboxymethyl-cellulose (with a MW of 1500 kDa) was dissolved indeionized water to the concentration of 2,2%, and crosslinking initiatedafter precooling to 3° C. by means of addition of adipic aciddihydrazide AAD (0.08%) and a small excess of the carbodiimide EDC(0.5%) and buffered to a pH-value of 5.6 by means of an excess of MESbuffer (54 mM). 20 mL of the reaction mixture was placed at −15° C. in aglass mold measuring 1 mm in depth and 16 cm diameter. After 20 hours,the scaffold was thawed and washed in 50 mL of DI water. The next stepconsisted in fracturing the scaffold. For this the bulk scaffold wasstuffed into a 50 mL syringe and extruded through a 20G needle byapplying a pressure of 15 bars. The fractioned material obtained wasfurther washed with 50 mL of a saline solution containing 0.45 g ofNaCl. After the washing step, the material was autoclaved in a bottle ofglass containing 90 mL of DI water using a temperature of 118° C. during24 minutes. The content was then put onto a filter device with a poresize of 0.22 um and fluid withdrawn by briefly applying a suctionpressure of 750 mbar such as to obtain a final volume of 10 mL Thematerial was then transferred into a syringe with luer lock forinjection.

Example 4

The fractioned material obtained in example 1 was further washed withphosphate buffered saline (PBS). The washing step was performed bythawing the fractioned material in a bath of saline solution. 10 mL ofthe fractioned material obtained in example 1 consisting of 0.6 g of drypolymer and of 9.4 g of water was washed with 50 mL of a saline solutioncontaining 0.45 g of NaCl.

After the washing step, the material was autoclaved in a bottle of glasscontaining 90 mL of DI water using a temperature of 121° C. during 20minutes. The content of the bottle was then centrifuged using anacceleration of 4 g during 2 minutes; 50 mL of water was removed using aBecher and a pipette to obtain the final consistency. The consistencywas adjusted by addition or withdrawal of fluid on a filter device; thefinal volume was about half of the original fabrication volume.

C) Comparative Tests

Example 5.1

The Young's modulus of the soft tissue engineering material according tothe invention, in conjunction with particle geometry and hydrationlevel, enables the deployment of the branched sheets of the particlesand consequently the 3D projection of the volume created (see FIG. 6 ).

When the Young's modulus is too low, or the hydration too large, thematerial does not project in 3D but spreads (picture A of FIG. 6 ). Whenthe Young's modulus and hydration level are correct, the materialcreates a 3D implant, stable over time (picture B of FIG. 6 ).

The effect of the mechanical properties of the soft tissue engineeringmaterial was further evaluated quantitatively by evaluation of theshort-term (3 weeks) implantation behavior as a function of themechanical properties of the implant. For this purpose, soft tissueengineering material fabricated according to 6 different recipes andcharacterized by their deployment pressure and Young's modulus of thesoft tissue engineering material. The materials were injectedsubcutaneously in mice, and the implant evaluated with regard toundesired spreading from the injection site, evolution of volume for thefirst hour and then at three weeks, as well as regarding stability ofshape and creation of a 3D projection. The results are summarized intable 1:

TABLE 1 Young's modulus (soft tissue Deployment engineering In-vivo 3DIn-vivo shape Recipe pressure material) deployment projectionmaintenance #1 4 Pa 40 Pa No No Flows #2 19 Pa 0.13 kPa No No Flows #332 Pa 0.28 kPa Inconsistent Inconsistent Inconsistent #4 95 Pa 0.74 kPaYes Inconsistent Inconsistent #5 163 Pa 1.5 kPa Yes Yes Yes #6 274 Pa3.3 kPa Yes Yes Yes, but too hard to the touch

They indicate for that for the implantation site and procedure chosen, aminimum of about 100 Pa of deployment pressure is needed to obtain adesired consistent (yet slight) volume swelling upon implantation, andthat a Young's modulus of at least 1.5 kPa is required for stable 3Dprojection (not surprisingly, this approximately matches the knownYoung's modulus of 2 kPa for adipose tissue). Only slightly higherYoung's moduli (3.3 kPa) are perceived as unnaturally hard to the touchfrom the outside. The Young moduli indicated are drained moduli; theundrained values are about 2.5×higher. The Poisson ratio under drainedconditions was near zero, whereas it was near 0.5 for undrainedconditions.

Uniaxial compression used for Young modulus determination wasessentially perfectly reversible to high strains (at least 30%), bothfrom geometric observation and return to baseline force within a fewpercent of the maximum force in particular for the drained conditions.

To further characterize the mechanics of the soft tissue engineeringmaterial, we analyzed samples obtained with recipe #5 of Table 1 inoscillatory rheology, and in uniaxial creep tests. For rheology, we useda HaakeRS100 RheoStress device, FL16 vane geometry with factorysettings, stress sweep from 1 Pa to 100 Pa at constant 1 Hz frequency.At low stress (<10 Pa), the sample behaves like an elastic solid withminor viscous contribution (elastic modulus G′ on the order of 5 kPa,viscous modulus G″ about 0.9 kPa), whereas at higher stresses (20-30 Paof shear stress in the FL16 vane geometry), a yield point is observedand the sample starts to flow with G′ approaching G″; however, as soonas the movement is stopped, the samples recover their original G′ and G″values at low frequency and stress (essentially perfect repeatability ofthe experiment without need for a setting period). This reversible, butnonlinear viscoelastic behavior contributes to injectability of thematerial (at shear stresses beyond yielding), and simultaneously itspropensity to rapidly regain its stable solid-like properties oncemovement ceases.

Creep addresses how a material behaves under a constant load. Weassessed creep during uniaxial compression (samples of an about 5 mmheight under a chuck of 5 cm diameter), and found an uncommon behavior:For all pressures safely accessible to the uniaxial compression machineused (<2.5 kPa), chuck movement would completely stop at a finite sampleheight, indicating that for slow compression, the samples can withstandvery substantial pressures equal to their Young modulus or higher. Localdensification due to particle compressibility as well as efficientparticle interlocking in the engineering material according to theinvention are at the origin of this particular behavior. Surprisinglythe effect is protective for the shape achieved in-vivo under slowlyapplied pressures (for instance, an individual lying down on an injectedsite).

Deployment at low pressures enables the gentle aspiration of tissues orcell suspensions for co-grafting applications (mixing with adiposetissues and injection of the mixture to create a living volume). FIG. 7shows the percentage of cells retained in the material obtained afterseeding fibroblast cells using the deployment effect or using only apassive seeding (no deployment effect).

To achieve cell adhesion for the experiment described in relation toFIG. 7 , material as prepared in example 3 was coated with collagen I(10% of the mass of CMC) in an sodium acetate buffer pH 4, followed DIrinsing and covalent attachment of the adsorbed collagen by use of EDC(10 mg/mL, in 100 mM MES buffer at pH 5.5), followed by inactivation ofremaining EDC in basic pH and readjustment to physiological pH with PBSbuffer. All steps made use of the deployment effect to enhance fluidexchange; the coating protocol is a result of optimization with respectto total collagen adsorption efficiency, homogeneity, collagen densitylining the pores, absence of fibril formation and cell adhesion.

For measuring the effect of the deployment advantage of the particlesaccording to the invention, two different materials were injected inmice:

-   -   one which was partially hydrated, and once injected, deployment        of the particles took place by taking up interstitial fluids;        and    -   another material the particles of which were fully hydrated, and        once injected, would not deploy itself because the channel-like        conduits were already “full” of fluid and therefore was not        capable to deploy more.

In both cases the percentage of the implant area occupied by cells andcollagen or other proteins (“cellularization”) was evaluated as shown inFIG. 8 .

Further experiments were conducted with the two materials in order toconfirm working hypothesis that the deployment of the partially hydratedparticles by means of their peripheral protrusions was producing azipper effect leading to stability of the shape of the injected material(implant) and of the volume created and preventing migration of theparticles in the body.

The results obtained with the material with deployment ability clearlyshowed its superiority as represented in Table 2:

TABLE 2 Height of the Height of the implant 3 days implant measuredStandard after the Standard after the injection deviation injectiondeviation Material (mm) (mm) (mm) (mm) With 3 1 3 1 deployment abilityWithout 2 1 0.5 1 deployment ability

Surprisingly it seems that the material with deployment ability isfrictioning with the surrounding tissues enabling the material to stayin place (anchoring effect).

Example 5.2

Since isotropicity of the conduits seems to play a major role in thedeployment capability of the material further experiments were conductedin this regard. Indeed particles with high channel anisotropicity havelong, highly oriented, parallel channels, and will collapse easily inthe direction perpendicular to the channel orientation and therefore beunable to deploy correctly. In cross-sections of the particles, thisanisotropy is visible by the occurrence of channels with very largeratios of longer to smaller diameter.

In order to verify these assumptions particles were manufactured withnon-isotropic conduits and used for the manufacture of an implantablesoft tissue engineering material comprising a multitude of suchparticles.

This material was compared to the material according to the invention bymeasuring the height of the implanted material immediately after theinjection into the body and after 3 days. The results are shown in thebelow table. It was observed that the 3D deployment was reduced in thenon-isotropic like conduits as shown in table 3:

TABLE 3 Height of the Height of the implant implant measured af-Standard measured 3 Standard ter the injection deviation days after thedeviation Material (mm) (mm) injection (mm) (mm) With isotropic 4 1 3 1conduits With non- 2 1 0.5 1 isotropic conduits

D) Role of the Mean Diameter of the Conduits on the Vessels Ingrowth

Since mean diameter of the conduits plays a major role in thevascularization of the material once implanted in vivo, furtherexperiments were conducted in this regard. The graph in FIG. 12 showsthe vessels density (number of vessels/mm2) observed on histologypictures after the implantation of a soft tissue engineering materialhaving a mean diameter of the conduits of 16 micrometers and a softtissue engineering material having a mean diameter of the conduits of127 micrometers. We observe that the vessels density is significantlylower in the case of the smallest mean diameter of the conduits.

E) Clinical Use of the Implantable Soft Tissue Engineering MaterialAccording to the Invention

Example 6.1

Prior the intervention, the surgeon using the soft tissue engineeringmaterial defines the areas where new volumes are needed. For this,he/she evaluates visually the volume defects and traces lines using amarker defining the future injection lines.

10 mL of the soft tissue engineering material was placed in a plasticsyringe equipped with a Luer-lock connector tip (corresponding to 0.6 gof dry mass of polymer). A cannula was connected to the tip and insertedin the target area of the patient through a thin skin incision. Oncepositioned in the target site, for example between the subcutaneousadipose layer and the pectoral muscle in a woman breast, the cannula iswithdrawn at a speed of 0.5 cm/s while 10 mL of the material is injectedby the surgeon by applying a pressure on the piston of the syringe of4000 N/m². The injection can be repeated in a neighboring area, enablingto increase the total volume injected.

Example 6.2

The injection of 10 mL of the soft tissue engineering material isrepeated by using the same incision point as in example 6.1. but bymodifying the direction and the angle of the cannula between eachinjection.

The surgeon performed one incision through the skin of the patient closeto the area needing volume enhancement. The Luer-lock syringe containingthe soft tissue engineering material was screed to a cannula (14G forexample) and the cannula was inserted in the patient's tissues. Thecannula was inserted into the tissues up to reaching the target and theinjection of 10 mL of the soft tissue engineering material was startedby applying a pressure of 4 kPa to the piston of the syringe whilewithdrawing the cannula in the direction of the incision point. Then,without taking the cannula out of the patient's body, the empty syringewas unscrewed and a new filled syringe containing the soft tissueengineering material was screwed on and a new injection was performed ina new direction of interest, predefined by marked lines on the patient'sskin.

Example 6.3

In another embodiment, the soft tissue reconstruction material is firstmixed with adipose tissues from the patient using two syringes and aconnector before being injected as a mixture into the target area usinga cannula.

Example 6.4

In one embodiment, the soft tissue engineering material is combined withthe graft of adipose tissues preliminary harvested from the patient. Forexample, adipose tissues are extracted by liposuction using a harvestingcannula connected to a 10 mL Luer-lock syringe. Tissues are let sedimentfor 5 minutes allowing to remove the blood and oil floating above theadipose tissues. In one embodiment, the user injects one spaghetti ofadipose tissues of 2 mL to 10 mL and then he/she injects a spaghetti ofthe soft tissue engineering material. In another embodiment, the adiposetissues are mixed with the soft tissue engineering material byconnecting two syringes (one containing the adipose tissues, the otherone containing the soft tissue engineering material) using a Luer-toLuer connector and by pushing sequentially on the two pistons of the twosyringes until obtaining a homogeneous mixture. The mixture obtained isthen injected using the injection method described before.

Example 6.5

In another embodiment, the material is injected in the target area andthe shape of the implant is shaped manually by the surgeon from theoutside of the patient in order to create the shape required.

Example 6.6

In one embodiment, the implantable soft tissue engineering material issterile and contained in a syringe. It is delivered in the target areaof the patient using a tubular element such as a sterile Luer-lockinfiltration Coleman cannula of 14 Gauge. Typically, the material isinjected into subcutaneous tissues, into adipose tissues, into musculartissues, between two layers of the above-mentioned tissues. For thedelivery, the user performs first a small incision (1 mm to 4 mm inlength) located at least at 2 cm of the targeted injection site. Theuser inserts the cannula through the incision up to reaching thetargeted point, located at 2 cm to 15 cm from the insertion point.He/she then injects retro-gradually 5 mL of the soft tissue engineeringmaterial by pushing gradually on the piston of the syringe whilewithdrawing the cannula from the targeted point to the incision point.So doing, the user injects a spaghetti like volume having a diametercomprised between 1 mm and 8 mm, enabling the integration of the softtissue engineering material within the surrounding tissues. Theprocedure can be repeated several times from the same injection point inorder to create a 3D arrangement of spaghettis. The localization of thespaghettis is controlled manually by the user, who is able to evaluatethe depth of the injection and the localization in the different planesof the patient's tissues.

Other Variations of Examples 6.1. To 6.6. Are Described Below

In one embodiment, the user uses his/her hands to press on the skin ofthe patient while inserting the cannula and injecting the material inorder to maintain the patient's tissues from the outside and to definethe localization of the material.

In one embodiment, the user injects the material using the same devicedescribed previously but injects the material in a bolus shape, which isexpanding the surrounding tissues of the injection site.

In one embodiment, the soft tissue engineering material is combined withthe graft of adipose tissues preliminary harvested from the patient. Forexample, adipose tissues are extracted using a harvesting cannula byliposuction. Tissues are let sediment for 5 minutes allowing to removethe blood and oil floating above the adipose tissues. The adiposetissues are distributed in 10 mL syringes. In one embodiment, the userinjects one spaghetti of adipose tissues of 2 mL to 10 mL using theColeman method and then he/she injects a spaghetti of the soft tissueengineering material. In another embodiment, the adipose tissues aremixed with the soft tissue engineering material by connecting twosyringes (one containing the adipose tissues, the other one containingthe soft tissue engineering material) using a Luer-to-Luer connector andby pushing sequentially on the two pistons of the two syringes. Themixture obtained is then injected using the method described before.

In another embodiment, the soft tissue engineering material is manuallydistributed in a body cavity (such as a breast cavity after siliconeimplant removal) using a sterile spatula in order to create a layer ofthe soft tissue engineering material.

In another embodiment, the soft tissue engineering material is suturedto surrounding tissues (in the case of large particles).

F) Clinical Results Obtained and Comparative Studies with Prior ArtMaterials

Example 7

A comparison of the stability and migration of 4 different materials,including the soft tissue engineering material according to theinvention was performed. The materials were the following:

“HA 1” is a commercially available hyaluronic acid based filler(“Juvederm Ultra 2” from Allergan.

“HA 2” is a commercially available, strongly crosslinked hyaluronic acidbased filler (“Macrolane” from Q-med AB).

“Matrix” is a commercially available, collagen based, flowable matrixused for wound repair (from Integra LifeSciences corporation).

“Material developed” is the material obtained in examples 1 to 4.

A defined volume of the different tested items (200 microliters) wasinjected subcutaneously in CD1 female mice in the back area of theanimal. Two samples of each tested item were injected, namely one oneach side of the spinal cord of the animal. In the case of the siliconeitem, the samples were implanted by first performing an incision in theskin of the animal and by inserting manually with tweezers the layer ofsilicone. The volumes of the items were monitored over time usingexternal measurements with a Caliper and using MRI scanning and MRIimages analysis. After 3 and 6 months, the animals were euthanized andhistology of the different implanted materials was performed.Bio-integration (percentage of the material occupied by cells andtissues, vascularization) was quantified. The results are represented inFIG. 9 , which shows macroscopic pictures of the implants site in micefor the different tested items, at different time points (t=0 is justafter the injection step). In the drawings on the left side of FIG. 9 ,the dashed lines represent the implant localization just after theinjection. The grey surface represents the implant localization 3 monthsafter the injection.

The macroscopic observation of the histology samples (see FIG. 10 )enabled to show that the material developed lead to the growth ofvessels and tissues, whereas the HA-based materials tested remainedclear and transparent, showing neither tissular nor vascular ingrowth.The

Matrix 1 material was populated with cells but it degraded before the 6months timepoint. These results are represented graphically in FIG. 11 .

The presence of a capsule surrounding the implants compared wasinvestigated on histological sections stained with Masson trichrome. Anadditional material was included in this comparative study, namely“Silicone” which a silicone layer sample cut from a silicone tissueexpander used in breast (Natrelle 133) Tissue expander from Allergan.The implanted samples were squares of 6 mm side and measured 1.5 mm inthickness.

The thickness of the capsule was measured for each material tested. Theresults are presented in Table 4 below:

TABLE 4 Material Material Matrix implanted developed HA1 HA2 1 SiliconeThickness of No No 92 +/− 17 No 104 +/− 20 the capsule capsule capsulemicrometers capsule micrometers (3 months after the implantation

It was observed that the soft tissue engineering material according tothe invention was stable over time. It did not migrate or increase involume. On the contrary, HA 1 was increasing in volume and the twoimplants merged together (they moved from the initial position). HA 2was also stable but the histological analysis showed the presence of aforeign body reaction (thin capsule around the implant, presence ofgiant cells), which could explain the stability of the position. Thematerial was isolated from the body and did not migrate. The matrix 1material was rapidly resorbed and did not produce a durable volume.

Although the invention has been described in conjunction with specificembodiments thereof, it is evident that many alternatives, modificationsand variations will be apparent to those skilled in the art.Accordingly, it is intended to embrace all such alternatives,modifications and variations that fall within the scope of the appendedclaims.

It is appreciated that certain features of the invention, which are, forclarity, described in the context of separate embodiments, may also beprovided in combination in a single embodiment. Conversely, variousfeatures of the invention, which are, for brevity, described in thecontext of a single embodiment, may also be provided separately or inany suitable subcombination or as suitable in any other describedembodiment of the invention. Certain features described in the contextof various embodiments are not to be considered essential features ofthose embodiments, unless the embodiment is inoperative without thoseelements.

Particle Suitable for the Manufacture of an Implantable Soft TissueEngineering Material

BACKGROUND OF THE INVENTION 1. Field of the Invention

The invention relates to a particle suitable for the manufacture of animplantable soft tissue engineering material, to an implantable softtissue engineering material comprising a multitude of particles andmethod for manufacturing such particles.

2. Description of the Related Art

From U.S. Pat. No. 7,993,679B2 a flowable collagen/glycosaminoglycanmaterial is known which can be used as a wound dressing. This materialhas no capacity to be expandable.

From U.S. Pat. No. 5,629,191 a method is known for making a mass ofgeneral spherical porous matrix particles, which, however, have nocapacity to undergo expansion upon uptake of liquid and have sphericalshapes with no protrusions.

From the scientific article “Injectable preformed scaffold withshape-memory properties”, Bencherif et al., Proceedings of the nationalacademy of sciences, vol. 109, no. 48, Nov. 12, 2012, pages 19590-19595a process is known in which a polymerizable material is cross-linkedusing a radical polymerization process starting at a temperature of +4°C. and continued at an unknown temperature. The premixed polymerizationmixture is then poured on a pre-cooled mold producing a skinning effectconsisting of the partial closing of external pores. The polymerizedscaffold material is then used as such for injection into a patient,i.e. with no further fractioning of the scaffold obtained.

Other prior art manufacturing processes start from pre-existing tissuesas a raw material for producing particulate implantable soft tissuematerials. The disadvantage of such methods lays in the fact thathardness and elasticity of the particles, which are essential for theintended volume reconstruction process, cannot be adjusted.

Other known processes start from frozen material, which gives lesselastic and less porous particles.

A) General Definitions

Soft Tissue Engineering Material

A biocompatible material for use in reconstruction, repair, enhancement,augmentation or substitution of soft tissues. In the context of thisinvention, a material obtained by combination of the branched-sheet 3Dparticles with a physiologically acceptable fluid (physiological saline,phosphate buffered saline, blood plasma, living aspirates such aslipoaspirate).

Dry Mass Concentration

The dry mass concentration of the soft tissue engineering material canbe determined by standard techniques known in the art, i.e. washing ofthe particles to remove salts followed by drying to constant weight, andrelating the dry weight found in this way to the volume of the softtissue engineering material.

B) Definitions Concerning the Reference Conditions

Reference liquid and reference measurement condition The geometric andmechanical properties of sheet material, particles, and soft tissueengineering material all depend on the suspension liquid. It istherefore necessary to specify the reference liquid in which such valuesare measured. For the present invention, this liquid is phosphatebuffered saline, with the following concentrations: potassium chloride(KCl) 2.6 mM, potassium phosphate monobasic (KH₂PO₄): 1.5 mM, sodiumchloride (NaCl) 138 mM, sodium phosphate dibasic (Na₂HPO₄): 8.1 mM indeionized water. Knowing these concentrations, a person skilled in theart will be able to prepare the reference liquid. The measurement ofgeometrical and mechanical properties of samples should be carried outwith samples immerged for at least 24 h with at least a 100× excess ofreference liquid relative to the sample volume prior to the measurementand immerged in the reference liquid during the measurement. Thereference temperature is room temperature (25° C.).

C) Geometric Definitions

Particle Size

The size of a particle is defined by the volume of its convex hull. Theconvex hull is a well-known mathematical concept; it comprises all thematerial points making up the particle, and also all possiblestraight-line connections between these material points. The convex hullcan be determined by confocal imaging of freely suspended, isolatedparticles in the reference liquid. A non-exclusive example ofdetermination of the size of a particle in accordance with the inventioninvolves the following steps, known to a person skilled in the art:

1) Confocal microscopy according to techniques known in the art is usedto acquire an image of an isolated particle freely suspended in thereference liquid. Confocal imaging can typically be done based onautofluorescence of the sheet material or fluorescence from dyesdisplaying affinity or on the contrary being excluded from the sheetmaterial. 2) The confocal images acquired are stored electronically asstacks of images, each image corresponding to a confocal plane. Thepixels making up these images are also referred to as voxels since theyrepresent the fluorescence information from a confocal volume of acertain depth.

3) Each voxel is then either assigned to the sheet material or the voidspace (based on thresholding between two very distinct populations ofintensities).

4) Using this binarized representation, the convex hull constructedusing available software. For the particles at hand, the convex hulldefines a mostly polyhedral 3D volume that contains all the voxels ofthe branched sheet but also the pore space interior to the particle.

5) Finally, the total particle volume can be calculated by the number ofvoxels identified as belonging to the convex hull times the volumeassociated with each voxel.

Equivalent Particle Diameter

The equivalent particle diameter is the diameter of sphere that has thesame volume as the particle's convex (the “particle size”). To bespecific, if the volume of the particle is V, then the equivalentdiameter is given d=(6*V/pi)∧(⅓), where pi=3.1419 . . .

Envelope of a Particle

The envelope of a particle is defined by the surface of its convex hull.As a non-exclusive example, a digital representation of the convex hullof a given particle can be obtained by the sequence outlined for theparticle size above (i.e. confocal images of freely suspended particlesin the reference fluid are acquired, stored, and analyzed as describedfor the size of the particle to define its convex hull). The envelope ofthe particle is defined by the voxels of the convex hull that have atleast one direct neighbor (sharing a face or corner) that does notbelong to the convex hull.

Protrusions of the Particle

The protrusions with the particle surface are defined as the parts ofthe sheet that reach the envelope of the particle as defined above. As anon-exclusive example, a digital representation of the protrusions canbe defined from the approach described for the envelope of the particle.Indeed, once the digital representation of the envelope is defined, thelocation of the waU protrusions is found by taking the intersection ofthe voxels representing the branched sheet and the voxels representingthe particle envelope.

Protrusion Depth

The protrusion depth is defined by the length of a normal vector from apoint of the particle envelope to the closest intersection point with asheet of the particle. If the point of interest of the particle envelopbelongs itself to a sheet, the associated protrusion depth is zero. Ifthe normal vector has no intersection with a sheet, the distance betweenthe two intersection points of the normal vector with the envelope istaken instead.

Maximal Protrusion Depth

The maximal protrusion depth is largest observed protrusion depth for agiven particle.

Relative Maximum Protrusion Depth

The relative maximum protrusion depth is defined as the maximumprotrusion depth divided by the equivalent particle diameter. If largerthan 1, a value of 1 is recorded.

Contact Area of the Protrusions

The contact area of the protrusions is the area where the distance fromthe envelope to the nearest sheet is less than the average sheetthickness TC.

Conduits

The conduits are the voids inside the particles. In the soft tissueengineering material, they can be filled with a physiologicallyacceptable fluid.

Conduit diameter D_(C)

The conduit diameter D_(C) is defined as the mean Feret diameter of theconduits as individual void areas on a section of the particle. Todetermine the conduit diameter D_(C), 2D cross section images ofparticles freely suspended in reference liquid may be obtained, forinstance by confocal imaging. On such sections, the conduits areidentified as the visible distinct voids; for conduits reaching theparticle exterior on the section of interest, the boundary of the convexhull of the particle as visible on the 2D section is used to completedelimitation of the void by the sheet material. For each of the voidsidentified, the Feret diameter, known to a person skilled in the art, ismeasured in all possible orientations (in practice, a few hundred), andthe average taken for each void. To obtain the average conduit diameter,the average of these averages is taken, ideally for several sections ofseveral particles.

Thickness of the Sheet T

The average thickness of the sheet T is defined as the mean thickness ofsheet material encountered when randomly slicing the particle. Todetermine the average thickness of the sheet T, 2D cross section imagesof particles freely suspended in reference liquid may be obtained, forinstance by confocal imaging. On such images, pixels are eitherattributed to the sheet material or the pore fraction, typically basedon intensity thresholding. To determine the sheet thickness, randomstraight lines are drawn across the particle (random orientation, randomlocalization). For each such line, transitions from void to sheet andback are noted, each such transition from void to sheet and backrepresenting a sheet intersection. The local thickness of the sheet forsuch a sheet intersection is determined from the number of pixelsbelonging to the walls and the known size of the pixels. The averagethickness of the sheet T is obtained by averaging over all the sheetintersection, ideally from hundreds of lines per particle and also astatistically relevant number of particles.

Relative Sheet Thickness

The relative sheet thickness is the ratio of the sheet thickness toconduit diameter: T_(relative)=T/(D_(C)+T). For particles with lowrelative sheet thickness T_(relative) (i.e. T_(relative)<0.1) oneexpects the total sheet volume fraction to be proportional toT_(relative), with a proportionality constant near unity.

Conduit Diameter, Sheet Thickness and Relative Sheet Thickness in theSoft Tissue Engineering Material

Conduit diameter, sheet thickness and relative sheet thickness are notonly defined for the isolated particles, but also for the soft tissueengineering material. In this case, they are understood to depend on theamount and nature of the fluid added to obtain the soft tissueengineering material. They can be measured by the same methodology asthe one exposed for the particles, i.e. by acquisition of planar imagesfor instance by confocal imaging, following by the image analysistechniques defined above. In the case of the soft tissue engineeringmaterial, one should be aware that the conduits detected as voids willgenerally be conduits within particles, but also conduits formed betweenneighboring particles; except for if particular precautions such aslabeling of different particles with different dyes are taken, it is notgenerally possible to distinguish between the two. Likewise, some of thesheet intersections in the procedure for determining sheet thicknesswill be formed by touching sheets from neighboring particles, withoutthis being necessarily evident in image analysis.

Branched Sheet Morphology

The branched sheet morphology applies to the particles. Porous materialsin general are classified either as “open” or “closed” foams (CellularSolids: Structure and Properties, Gibson and Ashby). In the terminologyof foam science, closed foams have “cells” (pores, voids) where thefaces shared with neighboring “cells” are solid membranes (CellularSolids: Structure and Properties, Gibson and Ashby); these closed foamsentrap the pore fluid, which cannot easily escape through the solidmembranes delimiting the “cells”. Classical open foams have no membranesbetween neighboring cells, but only struts along the edges where threeor more cells meet (Cellular Solids: Structure and Properties, Gibsonand Ashby). Such open foams naturally have a fully interconnected porespace, and a fibrous network of solid material. The branched sheetmorphology can be seen as a special structure that functionally combinesaspects of open and closed foams. In a perfect branched sheetmorphology, each “cell” can be understood to have both open and closedfaces. The open faces provide connections to neighboring “cells”, whilethe closed faces make up the bulk of the branched sheet. The open facesare organized in such a way to allow connectivity to the particleexterior for the vast majority of the foam cells, and hence to providepore interconnectivity. The closed faces together provide the branchedand warped 3D-sheet morphology. Typically, each “cell” has at least twoopen and two closed faces, where an open face can also be a connectionto the particle exterior.

Anisotropicity Index

The anisotropy index is the ratio between the longest Feret diameter tothe smallest Feret diameter of a body or void. In the context of thisinvention the anisotropicity index refers to the conduits, i.e. theratio between longest and smallest Feret diameter for each conduit, asindividualized as a distinct void in 2D cross-sectional images (seeconduit diameter D^(C)). The global anisotropy index is the average ofthe anisotropy indexes associated with the individual conduits, asidentified as individual voids on 2D cross-sections.

D) Definitions Concerning Mechanical Properties

Younq's Modulus

The Young's modulus of a given material is the ratio between stress(force per area) and strain (deformation relative to the originallength) as measured in uniaxial compression, typically for small strains(10% compression or less). For the present invention, Young's moduli aredefined at three different scales: the Young's modulus of the materialfrom which the branched sheets are made; the Young's modulus of theresulting particles; and the Young's modulus of the assembled softtissue engineering material.

Young's Modulus of the Sheet Material

The Young's modulus of the sheet material is the Young's modulus asmeasured on a pure, non-porous sample of the sheet material, in thereference liquid. In some embodiments, the sheet material can not onlybe synthesized as porous, branched sheet particles, but also assufficiently large bulk material. In this case, the most precise methodof evaluation of the Young's modulus of the sheet material is by directmechanical testing of a bulk sample of the sheet material (underreference conditions). In other embodiments, it is difficult to producethe sheet material in homogeneous bulk quantities; in this case, theYoung's modulus of the sheet material can still be evaluated bymicroindentation techniques, directly on structured materials, as forinstance described in Welzel et al., Adv Healthc Mater 2014, or Beduerat al., Advanced Healthcare materials 2015. 4(2): p. 301-12, using thereference conditions given above.

Young's Modulus of the Particles

As there is no technique known to directly assess the Young's modulus ofthe irregular, porous particles of the present invention, macroscopicsamples with regular geometries (typically cylinders of a few mm heightsand 1 or 2 cm diameter), but with identical pore structure and sheetmaterial are produced for mechanical testing. It is then possible to usestandard mechanical testing equipment to impose the desired uniaxialdeformation, and to measure the associated force to determine strain(deformation relative to original sample height) and stress (force percross-sectional area) to calculate the Young's modulus of the particlesunder reference conditions. As it will be known to a person skilled inthe art, the mechanical testing of porous samples heavily depends on theboundary conditions for the pore fluid. Under the reference conditions,the sample is immersed in reference fluid, and is therefore underso-called drained conditions. Nevertheless, to avoid undue influence ofthe drainage of the pore fluid on the measured Young's modulus, it isimportant to ensure that the fluidic drainage resistance of the samplebecomes negligible with respect to the solid's response. In practice, itwill be known to a person skilled in the art that the primary variablesto control in this respect are sample geometry and compression rate, andthat suitable compression conditions can be found for instance throughminimization of hysteresis or validation by stress relaxation tests.

Young's Modulus of the Soft Tissue Engineering Material

A third Young's modulus applies to the soft tissue engineering material.The soft tissue engineering material consists of a given amount ofparticles (specified for instance relative to the original amountfabricated or absolutely in terms of dry mass present), and a specifiedamount of reference fluid. For a wide range of fluid volume to particlemass, this mix forms a paste with a non-zero yield stress and strain(measurable through rheometry), meaning that the paste behaves like asolid for sufficiently small deformations. For such small deformations,the Young's modulus can be measured by uniaxial compression. The Young'smodulus naturally depends on the ratio of particles to fluid, thecomposition and geometry of the particles, as well as the composition ofthe fluid.

In practice, it should be noted that soft tissue engineering materialcannot simply be immerged into the reference fluid, undue dispersal ofthe particles would result. Instead, the drainage of reference porefluid during compression has to be carried out through a porousmembrane, with a pore size such that the reference fluid easily passes,but the particles are retained in the compression area. A person skilledin the art will easily be able to adapt existing setups to this purposeor identify a suitable commercial apparatus.

Deployment Pressure

The deployment pressure is the pressure with which the soft tissueengineering material of this invention expands, and therefore attractsliquid. Like the Young's modulus of the soft tissue engineeringmaterial, the deployment pressure depends on the exact nature of theparticles suspended, the exact composition of the fluid used for itsmeasurement, and the ratio of particles to fluid. The deploymentpressure of the soft tissue engineering material can be measured bycontacting the soft tissue engineering material with a mesh or membranefreely permeable to the reference fluid, but not the particles of thesoft tissue engineering material. The deployment pressure is then thedifference between the pressure recorded in the soft tissue engineeringmaterial and the pressure in pure reference fluid in equilibrium withthe soft tissue engineering material. A possible readout technique isobservation of hydrostatic height differences in a U-tube configurationwith the separation membrane between soft tissue engineering materialand the pure reference fluid; another possibility is the use of pressuregauges, one where only the reference fluid has access, the other exposedto the soft tissue engineering material.

Maximal Elongation of the Sheet Material

The maximal elongation of the sheet material is the maximum relativedeformation of the sheet material compatible with essentially purelyelastic deformation in the reference liquid: beyond this maximalelongation, the sheet material fails by either onset of plasticdeformation or brittle fracture.

Swelling Value of a Material

The swelling value is defined as the relative weight change of amaterial from dry to humid state. This parameter is a ration, namely(Humid weight−Dry weight)/Dry weight. The measurement of the swellingvalue is usually performed in physiological saline 0.9% NaCl.

BRIEF SUMMARY OF THE INVENTION

It is an object of the invention to provide a particle suitable for themanufacture of an of implantable soft tissue engineering material withthe capability of the particles to increase the size of its conduits byfluid uptake and therefore to increase the volume of the particles up toa predefinable and controllable volume.

This is achieved thanks to the particular properties of the branchedsheets composing the particle. The invention solves the posed problemwith a particle as disclosed and claimed herein and with an implantablesoft tissue engineering material as disclosed and claimed herein.

The advantage(s) of the particles according to the invention and of theimplantable soft tissue engineering material comprising such particlesis to be seen in the ability of the particles to absorb water or aqueoussolutions, like body fluids, thereby causing the three-dimensionallywarped and branched sheet to deploy so that its envelope will occupy alarger volume. This volume is predefined by the conduits volume.

Compared to other materials the particles according to the inventionhave shown the unique property that they are capable of deploying backgently, yet rapidly under in-vivo conditions. The reason for this is tobe seen in the deployment pressure required for partial swelling thatmatches the in vivo physiological deployment pressure. For example, toosoft porous materials flow and cannot sustain a three-dimensionalvolume. Since these known materials are too soft its pores (or conduits)are prevented to actively deploy. On the other hand, too hard materialsprovide a foreign body not matching the mechanical properties of thesoft target tissues and are creating a local mismatch of Young'smodulus, at the origin of a local inflammatory reaction and possiblyfibrosis.

The capacity of deployment of the branched sheet of the particleaccording to the invention combined with the interconnected porosity inthe form of channel-like conduits enables a better bio-integration, i.e.ingrowth of tissues and vessels into the particle compared to state ofthe art particle without the capability of deployment.

The injectable material allows the filling of tissue defects of smalland medium size and even defects larger than 50 cm³.

The protrusions of the particles—which produce an unexpectedzip-fastener effect—are an essential feature of the invention. Thisfeature enables cohesivity of the engineering material implanted intothe patient. The protrusions are a key feature enabling cohesion betweenthe particles and avoid migration/permanent deformation of theengineering material after its injection into the patient.

A further essential feature of the invention is the degree ofanisotropicity of the conduits enabling the volume of the engineeringmaterial to be maintained.

It has been found that if the conduits are too anisotropic, the sheetswill collapse and stick to each other, leading to spreading of theimplanted engineering material and preventing creation of a stablethree-dimensional volume.

The deployment of the branched sheets of the particles enables a betterand

rapid ingrowth of cells, tissues, blood vessels, lymph and extracellularmatrices and more stability of the shape and volume of the implantedengineering material.

The zipper-effect of the protrusions enables to maintain the injectedshape by avoiding migration of the particles.

Thanks to the combination of the above mentioned features and to theactive deployment, the engineering material is able to enter into africtional relationship with the surrounding tissues enabling theengineering material to stay in place (anchoring).

If the implanted material is too hard, this anchoring effect is replacedby a cutting effect and the tissue is damaged. If the material is toosoft, the material is not anchoring and slides over the tissues.

Deployment at low pressures enables the gentle aspiration of tissues orcell suspensions for co-grafting applications, for example mixing of theengineering material with adipose tissues and injection of the mixtureto create a living volume.

The low deployment pressure results from particle design. Theapplication imposes the Young's modulus of the particles, but dependingon the particle structure, the deployment pressure can vary. To obtain alow deployment pressure for a given Young's modulus, it is essentialthat compression of the particles can occur as much as possible byevacuation of pore fluid, and not by compression of the incompressiblepore fluid or nearly incompressible sheet material itself, as thisnecessitates greater isotropic pressure. For this, the branched sheetmorphology is key: the pore fluid can easily be evacuated throughinterconnectivity, whereas the closed faces provide mechanical solidity,allowing to keep the sheet volume fraction low and avoidingsheet-to-sheet contact and compression. This combination gives thedesired low deployment pressure for the desired Young's moduli.

In a special embodiment of the invention the volume V_(E) of theenvelope is larger than 5-10⁻⁴ mm³, preferably larger than 4-10⁻³ mm³and most preferably larger than 0.03 mm³ The volume V_(E) of theenvelope is purposefully smaller than 4-10³ mm³, preferably smaller than260 mm³, and most preferably smaller than 30 mm³.

The thickness T of the sheet 2 is purposefully larger than 1 μm,preferably larger than 5 μm, The thickness T is purposefully smallerthan 1000 μm, preferably smaller than 100 μm, and most preferablysmaller than 20 μm. Surprisingly it has been found that a sheetthickness in that range allows obtaining a material which is transparentwhen imaged with ultrasounds imaging devices (such as the ones used inpatients.

The total surface of the sheet 2 is purposefully larger than 8-10 ⁻⁴mm², preferably larger than 1-10⁻² mm², and most preferably larger than10⁻¹ mm². The total surface of the sheet 2 is purposefully smaller than50000 mm², preferably smaller than 10000 mm².

The total volume V_(S) of the sheet 2 is purposefully larger than 8-10⁻⁷mm³, preferably larger than 5.10-6 mm³, and most preferably larger than5-10⁻⁴ mm³. The total volume V_(S) of the sheet 2 is purposefullysmaller than 50000 mm³, preferably smaller than 5000 mm³, and mostpreferably smaller than 200 mm³.

In a special embodiment the anisotropicity index of the conduits 5 islarger than 1.05, preferably larger 1.10. The anisotropicity index ispurposefully smaller than 10.0, preferably smaller than 5.0 and mostpreferably smaller than 3.0.

The material for the sheet can be chosen from non-fouling substancespoly-ethyleneglycol (PEG), poly-acrylamide,poly-(Hydroxyethyl)methacrylate, preferably from polysaccharides such ascellulose, methylcellulose, carboxymethylcellulose, agarose, polysucroseor dextran.

The sheet may also comprise a material chosen from the following group:carbohydrates, hydrogels, collagens, gelatins, peptides or extracellularmatrices. These materials confer biocompatibility and biodegradability.

The carbohydrate can be

(i) a polysaccharide, preferably a negatively charged polysaccharide;

(ii) an alginate;

(iii) hyaluronic acid; or

(iv) a carboxymethyicelluose.

These materials confer elasticity to the particle enabling a reversiblecompression capability and a reversible fluid intake capability Thesheet may also comprise a synthetic polymer, preferably chosen from thefollowing groups

(i) silicones;

(ii) polyurethanes;

(iii) polyolefins;

(iv) acrylates, preferably poly-acrylamide or poly-acrylic acid;poly-(Hydroxyethyl)-methacrylate and copolymers thereof;

(v) polyesters;

(vi) polyamides; or

(vii) polyimides.

The material of the sheet 2 has preferably a maximal elongation of VS/VEor more. This confers a reversible fluid intake capability to theparticle.

The material of the sheet 2 has preferably a maximal elongation of3T/D_(C) or more.

The material of the sheet 2 has purposefully a molecular weight in therange of 50 Da-10 MDa.

In a further embodiment the ratio of D^(C)/T is larger than 1.0,preferably larger than 2. Purposefully the ratio of D^(C)/T is largerthan 5, preferably larger than 7. The ratio of D^(C)/T may be smallerthan 500, preferably smaller than 100. Purposefully the ratio of D^(C)/Tis smaller than 50, preferably smaller than 30. The ratio of D^(C)/T maybe smaller than 29, preferably smaller than 15. The specific choice ofthe ratio of D^(C)/T combined with the choice of material enables thereversible compressibility of the particle and obtaining suitableparticle Young moduli.

In a further embodiment the mean diameter D_(C) of the conduits islarger than 1 micrometer, preferably larger than 14 micrometer.Purposefully the mean diameter D^(C) of the conduits is larger than 20micrometers, preferably larger than 50 micrometers. The minimumdimensions for the conduits enable cellular and vascular ingrowth intothe particle.

In a further embodiment the mean diameter D_(C) of the conduits issmaller than 10 mm, preferably smaller than 4 mm. The mean diameterD^(C) of the conduits may by smaller than 2 mm, preferably smaller than1 mm.

Purposefully the mean diameter D_(C) of the conduits is smaller than 600micrometer preferably smaller than 300 micrometers. The maximumdimension for the conduits enables a mechanical stability of thescaffold and tight contacts between the scaffold and the tissues. If thevalue for D^(C) is too high, the cellularization and vascularization isnot as high as if D_(C) is optimal.

It was surprisingly found that the choice of these parameters allowsobtaining a material which matches the speckled appearance of nativetissue when imaged with ultrasound imaging devices (such as the onesused in patients.

In a further embodiment the protrusions 4 have a mean relative maximumprotrusion depth in the range between 0.05 and 1.0, preferentiallybetween 0.15 and 0.8.

The shape of the three-dimensionally warped and branched sheet 2 ispurposefully flexible and in particular it is preferably reversiblyexpandable upon absorption or removal of a liquid by the biocompatiblematerial. Liquids suitable for absorption are water, aqueous solutions,blood or other body fluids.

In a further embodiment the Young's modulus of the biocompatiblematerial of the sheet 2 is at least 200 Pa, preferably at least 1′000Pa. Purposefully the Young's modulus of the biocompatible material ofthe sheet 2 is smaller than 500 kPa, preferably smaller than 50 kPa. Thematerial constituting the sheets should have a Young's modulus highenough to avoid compression of the particle at pressures that can befound in the body (fluid pressure of the interstitial fluid). If thematerial is too “soft”, it has become apparent that the sheets collapseone against to each-other and they are not able to expand back to obtaina certain “memory shape” effect observed with stiffer sheets

In a further embodiment the mean diameter D_(P) of the particle islarger than 2 micrometers, preferably larger than 10 micrometers. Thisfeature enables the particle to create a cohesive implant. If particlesare too small, there is the risk of migration of the particle in lymphor blood vessels, or phagocytosis by macrophages for example.

The mean diameter D_(P) of the particle is purposefully smaller than 5mm, preferably smaller than 2 mm. This limitation of the particledimension enables at the same time structurability and shapeability of apaste created by a plurality of particles and a cohesivity of the paste.

In a further embodiment the particle comprises at least 5 conduits,preferably at least 10 conduits. A small number of conduits in theparticle leads to mechanical properties which are less optimal and theparticle will have more chances to collapse.

In a further embodiment the ratio between the mean diameter of theconduit and the mean diameter of the particle D^(C)/D_(P) is larger than1.5, preferably larger than 2.0 Surprisingly it has been found that anoptimal D^(C)/D_(P) ratio avoids reduced porosity which would reducealso the tissue ingrowth capability. The optimal ratio also leads to afavorable the biodegradation time. Purposefully the ratio between themean diameter of the conduit and the mean diameter of the particleD^(C)/D_(P) is smaller than 20, preferably smaller than 10.

In a further embodiment the contact angle between water and thebiocompatible material of the sheet 2 is in the range of 0° to 90°,preferably in the range of 0°-60°. This feature allows an optimal liquidintake in the conduits of the particle.

Preferably the Sheet 2 is Reversibly Compressible.

In a further embodiment the particle is hydrated and preferablycomprises at least 0.05 weight-% of the biocompatible material based onthe total weight of the hydrated particle. Purposefully the hydratedparticle comprises at least 0.1 weight-%, preferably at least 0.5weight-% of the biocompatible material.

Purposefully the particle is hydrated and comprises at most 15 weight-%of the biocompatible material based on the total weight of the hydratedparticle. The hydrated particle may comprise at most 5 weight-%,preferably at most 3 weight-% of the biocompatible material.

The particle according to the invention may comprise severalthree-dimensionally warped and branched sheets 2.

The invention therefore is also directed to a composition comprising:

a) a multitude of particles according to the invention; and

b) a physiologically acceptable fluid.

The amount of fluid may be such that the particles are only partiallyhydrated. This has the advantage that the particles of the compositionretain still the ability to deploy by up-take of body fluids afterinjection into the body of the patient.

The invention is further directed to an implantable soft tissueengineering material comprising a multitude of particles according tothe invention, preferably in form of a malleable paste. The multitude ofparticles may be admixed with water or an aqueous solution or blood toform a malleable paste. The multitude of particles may also be admixedto adipose tissue. This has the advantage that the mixing of the pastewith adipose tissues allows creating a new volume of adipose tissues.

The implantable soft tissue engineering material according to theinvention is purposefully reversibly compressible after injection into apatient by uptaking liquid from the surrounding tissues. The uptake ofliquid occurs up to a fixed predefined amount corresponding to the fullexpansion state of the particles, i.e. until the conduits are filledwith liquid. Once the conduits are full, the expansion cannot gofurther. In vivo, there is then an equilibrium between the particlespressure and the pressure of the in vivo interstitial pressure. At thisstage the material is not deformable anymore guaranteeing that thevolume created for the patient cannot be deformed anymore.

The implantable soft tissue engineering material according to theinvention purposefully

exhibits a non-linear compression behavior and a Young's moduluscomprised between 100 Pa and 15 kPa.

The implantable soft tissue engineering material according to theinvention may be used as a shapeable tissue or organ body implant. Theimplantable soft tissue engineering material according to the inventionmay also be used for treating tissue defects, in particular tissuedefects caused by severe trauma or cancer ablation. I may also be usedfor breast reconstruction and for lipofilling. Further the material maybe used for aesthetic restorations in the face and the body.

The invention is further directed to a method for manufacturingparticles according to the invention comprising the following steps:

a) pre-cooling a polymerizable biocompatible material in an aqueoussolution at a temperature below 10° C.;

b) cross-linking the pre-cooled mixture at a temperature below 0° C.,preferably below minus 1° C.; and

c) fractioning the cross-linked biocompatible material obtained.

The cross-linking according to step b) may purposefully be performed ata pH-value of minimum 5.0, preferably minimum 5.5.

The cross-linking according to step b) may purposefully be performed ata pH-value of maximum 8.5, preferably maximum 7.5. The cross-linkingaccording to step b) may be performed at a pH-value of maximum 6.9,preferably maximum 6.5.

The cross-linking process in step b) should not be one based on aradical polymerization since it has been found that radicalpolymerization can pose a hazard problem because of incorporation ofinitiator and possible depolymerization during sterilization in vitro.

The cross-linker used in step b) may be adipic dihydrazide.

In a special embodiment the cooling process in step b) consists of twosub-steps:

-   -   (i) a first sub-step to a temperature in the range of 0° C. to        −15° C., preferably in the range of −2° C. to −12° C., followed        by a hold time for temperature equilibrium; and    -   (ii) a second sub-step to a temperature in the range of −80° C.        to −2° C.

The polymerizable biocompatible material has preferably a molecularweight of 50 Da-10 MDa.

A BRIEF DESCRIPTION OF THE DRAWINGS

A special embodiment of the invention will be described in the followingby way of example and with reference to the accompanying drawings inwhich:

FIG. 1 illustrates a three-dimensional view of an embodiment of theparticle according to the invention and its virtual envelope;

FIG. 2 shows the same view as FIG. 1 highlighting the protrusions at theperiphery of the particle by means of broken-line circles;

FIG. 3 shows the same view as FIG. 1 in which the thickness T of thesheet at various locations is indicated by arrows.

FIG. 4 shows the same view as FIG. 1 in which the thickness diameter DCof some of the conduits is indicated by arrows.

FIG. 5 shows the same view as FIG. 1 in which the protrusion depth isindicated by the length of a normal vector Nv from a point of theparticle envelope to the closest intersection point with a sheet of theparticle.

FIG. 6 shows 3-D projections of a material having a Young's modulus tolow (A) and of a material according to the invention (B).

FIG. 7 is a graphical representation of the percentage of cells retainedin the soft tissue engineering material according to the invention whenseed in vitro with and without deployment effect.

FIG. 8 is a graphical representation of the percentage ofcellularization of the soft tissue engineering material according to theinvention.

FIG. 9 shows macroscopic pictures of the implants site over time.

FIG. 10 represents photographs regarding histology and macroscopicobservation of bio-integration of comparative examples.

FIG. 11 is a graph showing cellular invasion and vascularization ofcomparative examples.

FIG. 12 is a graph showing the vessels density (number of vessels/mm²)observed on histology pictures after the implantation of a soft tissueengineering material having a mean diameter of the conduits of 16micrometers and a soft tissue engineering material having a meandiameter of the conduits of 127 micrometers.

DETAILED DESCRIPTION OF THE INVENTION

The following examples clarify the invention further in more detail.

A) Manufacture of the Particles

Example 1

Carboxymethyl-cellulose (with a MW of 700 kDa) was dissolved indeionized water to the concentration of 2%, and crosslinking initiatedafter precooling to 4° C. by means of addition of adipic aciddihydrazide AAD (0.07%) and a small excess of the carbodiimide EDC(0.4%) and buffered to a pH-value of 5.5. by means of an excess of2-(N-morpholino)ethanesulfonic acid (MES) buffer (50 mM).

The reaction mixture was placed at −20° C. in a mold. After 1 day, thescaffolds were thawed and washed in de-ionized water (DI).

The next step consisted in fractioning the scaffold. For this, a bulkscaffold or a bulk scaffold piece was placed in a plastic bag andcompressed and sheared manually to create the particles according to theinvention. In another embodiment, the bulk scaffold was extruded througha thin tubular element by applying a known pressure to obtain afragmented material.

The particle size was controlled by the pressure applied on the pistonof the syringe and by the size of the extruding cannula. Typically, apressure of 15 bars and a cannula of 14G was used.

Example 2a

The same procedure as in example 1 was followed but prior to freezing,the reaction mixture was distributed into a silicone mold using apipette of 10 mL. The silicone mold contained microstructuredstar-shaped cavities measuring 100 micrometers in diameter and 20micrometers in depth. The silicone mold was covered with a flatpolypropylene counterpart, squeezing excess liquid from the mold. Theassembly was then placed into a freezer at −20° C.

Alternative Methods for the Manufacture of Particles:

Particles were manufactured by placing the scaffolds into a mixer andmixing them.

Particles were manufactured by ink-jet printing, 3D printing, andadditive manufacturing.

Particles were manufactured by mixing the reaction mixture with aphotosensitizer (typically acrylamide monomer andN,N′methylenebis(acrylamide), freezing at −20° C. and photopolymerizingusing a UV lamp or a visible lamp.

Particles were manufactured by grinding a preliminary manufacturedscaffold during at least 30 s, for example using a mixing robot (forexample Kenwood Major Titanium KMM060).

Particles were manufactured by cutting and/or slicing a preliminarymanufactured scaffold using cutting blades, possibly organized innetworks.

It is important to note that a classical emulsion polymerization methodwould give nearly perfectly round particles and therefore would not leadto the desired structure of the particles according to the inventionwith significant protrusions.

Example 2b

A solution of 5% of hyaluronic acid monomers with a molecular weight of90 kDa, MES buffer pH6, adipic acid dihydrazide (2 mg/mL) was mixed withEDC (4 mg/mL) and poured onto a consolidated paraffin microspheresscaffold. The paraffin beads were prepared according to “Microspheresleaching for scaffold porosity control”, Draghi et al, Journal ofMaterial sciences: Materials in medicine, 16 (2005) 1993-1997. Themixture was incubated at room temperature during 24 hours after whichthe paraffin beads were dissolved by an excess of hexane. The obtainedscaffold was then rinsed with isopropanol, and a mix of isopropanol andwater (40%:60%) and followed by a rinsing step with water.

The obtained scaffold was then fragmented by applying an extrusion forceon the scaffold through a narrow tubular element.

B) Manufacture of an Implantable Soft Tissue Engineering MaterialComprising a Multitude of Particles According to the Invention

Example 3

Carboxymethyl-cellulose (with a MW of 1500 kDa) was dissolved indeionized water to the concentration of 2,2%, and crosslinking initiatedafter precooling to 3° C. by means of addition of adipic aciddihydrazide AAD (0.08%) and a small excess of the carbodiimide EDC(0.5%) and buffered to a pH-value of 5.6 by means of an excess of MESbuffer (54 mM). 20 mL of the reaction mixture was placed at −15° C. in aglass mold measuring 1 mm in depth and 16 cm diameter. After 20 hours,the scaffold was thawed and washed in 50 mL of DI water. The next stepconsisted in fracturing the scaffold. For this the bulk scaffold wasstuffed into a 50 mL syringe and extruded through a 20G needle byapplying a pressure of 15 bars. The fractioned material obtained wasfurther washed with 50 mL of a saline solution containing 0.45 g ofNaCl. After the washing step, the material was autoclaved in a bottle ofglass containing 90 mL of DI water using a temperature of 118° C. during24 minutes. The content was then put onto a filter device with a poresize of 0.22 um and fluid withdrawn by briefly applying a suctionpressure of 750 mbar such as to obtain a final volume of 10 mL Thematerial was then transferred into a syringe with luer lock forinjection.

Example 4

The fractioned material obtained in example 1 was further washed withphosphate buffered saline (PBS). The washing step was performed bythawing the fractioned material in a bath of saline solution. 10 mL ofthe fractioned material obtained in example 1 consisting of 0.6 g of drypolymer and of 9.4 g of water was washed with 50 mL of a saline solutioncontaining 0.45 g of NaCl. After the washing step, the material wasautoclaved in a bottle of glass containing 90 mL of DI water using atemperature of 121° C. during 20 minutes. The content of the bottle wasthen centrifuged using an acceleration of 4 g during 2 minutes; 50 mL ofwater was removed using a Becher and a pipette to obtain the finalconsistency. The consistency was adjusted by addition or withdrawal offluid on a filter device; the final volume was about half of theoriginal fabrication volume.

C) Comparative Tests

Example 5.1

The Young's modulus of the soft tissue engineering material according tothe invention, in conjunction with particle geometry and hydrationlevel, enables the deployment of the branched sheets of the particlesand consequently the 3D projection of the volume created (see FIG. 6 ).When the Young's modulus is too low, or the hydration too large, thematerial does not project in 3D but spreads (picture A of FIG. 6 ). Whenthe Young's modulus and hydration level are correct, the materialcreates a 3D implant, stable over time (picture B of FIG. 6 ).

The effect of the mechanical properties of the soft tissue engineeringmaterial was further evaluated quantitatively by evaluation of theshort-term (3 weeks) implantation behavior as a function of themechanical properties of the implant. For this purpose, soft tissueengineering material fabricated according to 6 different recipes andcharacterized by their deployment pressure and Young's modulus of thesoft tissue engineering material. The materials were injectedsubcutaneously in mice, and the implant evaluated with regard toundesired spreading from the injection site, evolution of volume for thefirst hour and then at three weeks, as well as regarding stability ofshape and creation of a 3D projection. The results are summarized intable 1:

TABLE 1 Young's modulus (soft tissue Deployment engineering In-vivo 3DIn-vivo shape Recipe pressure material) deployment projectionmaintenance #1 4 Pa 40 Pa No No Flows #2 19 Pa 0.13 kPa No No Flows #332 Pa 0.28 kPa Inconsistent Inconsistent Inconsistent #4 95 Pa 0.74 kPaYes Inconsistent Inconsistent #5 163 Pa 1.5 kPa Yes Yes Yes #6 274 Pa3.3 kPa Yes Yes Yes, but too hard to the touch

They indicate for that for the implantation site and procedure chosen, aminimum of about 100 Pa of deployment pressure is needed to obtain adesired consistent (yet slight) volume swelling upon implantation, andthat a Young's modulus of at least 1.5 kPa is required for stable 3Dprojection (not surprisingly, this approximately matches the knownYoung's modulus of 2 kPa for adipose tissue). Only slightly higherYoung's moduli (3.3 kPa) are perceived as unnaturally hard to the touchfrom the outside. The Young moduli indicated are drained moduli; theundrained values are about 2.5×higher. The Poisson ratio under drainedconditions was near zero, whereas it was near 0.5 for undrainedconditions.

Uniaxial compression used for Young modulus determination wasessentially perfectly reversible to high strains (at least 30%), bothfrom geometric observation and return to baseline force within a fewpercent of the maximum force in particular for the drained conditions.

To further characterize the mechanics of the soft tissue engineeringmaterial, we analyzed samples obtained with recipe #5 of Table 1 inoscillatory rheology, and in uniaxial creep tests. For rheology, we useda HaakeRS100 RheoStress device, FL16 vane geometry with factorysettings, stress sweep from 1 Pa to 100 Pa at constant 1 Hz frequency.At low stress (<10 Pa), the sample behaves like an elastic solid withminor viscous contribution (elastic modulus G′ on the order of 5 kPa,viscous modulus G″ about 0.9 kPa), whereas at higher stresses (20-30 Paof shear stress in the FL16 vane geometry), a yield point is observedand the sample starts to flow with G′ approaching G″; however, as soonas the movement is stopped, the samples recover their original G′ and G″values at low frequency and stress (essentially perfect repeatability ofthe experiment without need for a setting period). This reversible, butnonlinear viscoelastic behavior contributes to injectability of thematerial (at shear stresses beyond yielding), and simultaneously itspropensity to rapidly regain its stable solid-like properties oncemovement ceases.

Creep addresses how a material behaves under a constant load. Weassessed creep during uniaxial compression (samples of an about 5 mmheight under a chuck of 5 cm diameter), and found an uncommon behavior:For all pressures safely accessible to the uniaxial compression machineused (<2.5 kPa), chuck movement would completely stop at a finite sampleheight, indicating that for slow compression, the samples can withstandvery substantial pressures equal to their Young modulus or higher. Localdensification due to particle compressibility as well as efficientparticle interlocking in the engineering material according to theinvention are at the origin of this particular behavior. Surprisinglythe effect is protective for the shape achieved in-vivo under slowlyapplied pressures (for instance, an individual lying down on an injectedsite).

Deployment at low pressures enables the gentle aspiration of tissues orcell suspensions for co-grafting applications (mixing with adiposetissues and injection of the mixture to create a living volume). FIG. 7shows the percentage of cells retained in the material obtained afterseeding fibroblast cells using the deployment effect or using only apassive seeding (no deployment effect).

To achieve cell adhesion for the experiment described in relation toFIG. 7 , material as prepared in example 3 was coated with collagen I(10% of the mass of CMC) in an sodium acetate buffer pH 4, followed DIrinsing and covalent attachment of the adsorbed collagen by use of EDC(10 mg/mL, in 100 mM MES buffer at pH 5.5), followed by inactivation ofremaining EDC in basic pH and readjustment to physiological pH with PBSbuffer. All steps made use of the deployment effect to enhance fluidexchange; the coating protocol is a result of optimization with respectto total collagen adsorption efficiency, homogeneity, collagen densitylining the pores, absence of fibril formation and cell adhesion.

For measuring the effect of the deployment advantage of the particlesaccording to the invention, two different materials were injected inmice:

-   -   one which was partially hydrated, and once injected, deployment        of the particles took place by taking up interstitial fluids;        and    -   another material the particles of which were fully hydrated, and        once injected, would not deploy itself because the channel-like        conduits were already “full” of fluid and therefore was not        capable to deploy more.

In both cases the percentage of the implant area occupied by cells andcollagen or other proteins (“cellularization”) was evaluated as shown inFIG. 8 .

Further experiments were conducted with the two materials in order toconfirm working hypothesis that the deployment of the partially hydratedparticles by means of their peripheral protrusions was producing azipper effect leading to stability of the shape of the injected material(implant) and of the volume created and preventing migration of theparticles in the body.

The results obtained with the material with deployment ability clearlyshowed its superiority as represented in Table 2:

TABLE 2 Height of the Height of the implant 3 days implant measuredStandard after the Standard after the injection deviation injectiondeviation Material (mm) (mm) (mm) (mm) With 3 1 3 1 deployment abilityWithout 2 1 0.5 1 deployment ability

Surprisingly it seems that the material with deployment ability isfrictioning with the surrounding tissues enabling the material to stayin place (anchoring effect).

Example 5.2

Since isotropicity of the conduits seems to play a major role in thedeployment capability of the material further experiments were conductedin this regard. Indeed particles with high channel anisotropicity havelong, highly oriented, parallel channels, and will collapse easily inthe direction perpendicular to the channel orientation and therefore beunable to deploy correctly. In cross-sections of the particles, thisanisotropy is visible by the occurrence of channels with very largeratios of longer to smaller diameter.

In order to verify these assumptions particles were manufactured withnon-isotropic conduits and used for the manufacture of an implantablesoft tissue engineering material comprising a multitude of suchparticles.

This material was compared to the material according to the invention bymeasuring the height of the implanted material immediately after theinjection into the body and after 3 days. The results are shown in thebelow table. It was observed that the 3D deployment was reduced in thenon-isotropic like conduits as shown in table 3:

TABLE 3 Height of the Height of the implant implant measured af-Standard measured 3 Standard ter the injection deviation days after thedeviation Material (mm) (mm) injection (mm) (mm) With isotropic 4 1 3 1conduits With non- 2 1 0.5 1 isotropic conduits

D) Role of the Mean Diameter of the Conduits on the Vessels Ingrowth

Since mean diameter of the conduits plays a major role in thevascularization of the material once implanted in vivo, furtherexperiments were conducted in this regard. The graph in FIG. 12 showsthe vessels density (number of vessels/mm2) observed on histologypictures after the implantation of a soft tissue engineering materialhaving a mean diameter of the conduits of 16 micrometers and a softtissue engineering material having a mean diameter of the conduits of127 micrometers. We observe that the vessels density is significantlylower in the case of the smallest mean diameter of the conduits.

E) Clinical Use of the Implantable Soft Tissue Engineering MaterialAccording to the Invention

Example 6.1

Prior the intervention, the surgeon using the soft tissue engineeringmaterial defines the areas where new volumes are needed. For this,he/she evaluates visually the volume defects and traces lines using amarker defining the future injection lines.

10 mL of the soft tissue engineering material was placed in a plasticsyringe equipped with a Luer-lock connector tip (corresponding to 0.6 gof dry mass of polymer). A cannula was connected to the tip and insertedin the target area of the patient through a thin skin incision. Oncepositioned in the target site, for example between the subcutaneousadipose layer and the pectoral muscle in a woman breast, the cannula iswithdrawn at a speed of 0.5 cm/s while 10 mL of the material is injectedby the surgeon by applying a pressure on the piston of the syringe of4000 N/m². The injection can be repeated in a neighboring area, enablingto increase the total volume injected.

Example 6.2

The injection of 10 mL of the soft tissue engineering material isrepeated by using the same incision point as in example 6.1. but bymodifying the direction and the angle of the cannula between eachinjection.

The surgeon performed one incision through the skin of the patient closeto the area needing volume enhancement. The Luer-lock syringe containingthe soft tissue engineering material was screed to a cannula (14G forexample) and the cannula was inserted in the patient's tissues. Thecannula was inserted into the tissues up to reaching the target and theinjection of 10 mL of the soft tissue engineering material was startedby applying a pressure of 4 kPa to the piston of the syringe whilewithdrawing the cannula in the direction of the incision point. Then,without taking the cannula out of the patient's body, the empty syringewas unscrewed and a new filled syringe containing the soft tissueengineering material was screwed on and a new injection was performed ina new direction of interest, predefined by marked lines on the patient'sskin.

Example 6.3

In another embodiment, the soft tissue reconstruction material is firstmixed with adipose tissues from the patient using two syringes and aconnector before being injected as a mixture into the target area usinga cannula.

Example 6.4

In one embodiment, the soft tissue engineering material is combined withthe graft of adipose tissues preliminary harvested from the patient. Forexample, adipose tissues are extracted by liposuction using a harvestingcannula connected to a 10 mL Luer-lock syringe. Tissues are let sedimentfor 5 minutes allowing to remove the blood and oil floating above theadipose tissues. In one embodiment, the user injects one spaghetti ofadipose tissues of 2 mL to 10 mL and then he/she injects a spaghetti ofthe soft tissue engineering material. In another embodiment, the adiposetissues are mixed with the soft tissue engineering material byconnecting two syringes (one containing the adipose tissues, the otherone containing the soft tissue engineering material) using a Luer-toLuer connector and by pushing sequentially on the two pistons of the twosyringes until obtaining a homogeneous mixture. The mixture obtained isthen injected using the injection method described before.

Example 6.5

In another embodiment, the material is injected in the target area andthe shape of the implant is shaped manually by the surgeon from theoutside of the patient in order to create the shape required.

Example 6.6

In one embodiment, the implantable soft tissue engineering material issterile and contained in a syringe. It is delivered in the target areaof the patient using a tubular element such as a sterile Luer-lockinfiltration Coleman cannula of 14 Gauge. Typically, the material isinjected into subcutaneous tissues, into adipose tissues, into musculartissues, between two layers of the above-mentioned tissues. For thedelivery, the user performs first a small incision (1 mm to 4 mm inlength) located at least at 2 cm of the targeted injection site. Theuser inserts the cannula through the incision up to reaching thetargeted point, located at 2 cm to 15 cm from the insertion point.He/she then injects retro-gradually 5 mL of the soft tissue engineeringmaterial by pushing gradually on the piston of the syringe whilewithdrawing the cannula from the targeted point to the incision point.So doing, the user injects a spaghetti like volume having a diametercomprised between 1 mm and 8 mm, enabling the integration of the softtissue engineering material within the surrounding tissues. Theprocedure can be repeated several times from the same injection point inorder to create a 3D arrangement of spaghettis. The localization of thespaghettis is controlled manually by the user, who is able to evaluatethe depth of the injection and the localization in the different planesof the patient's tissues.

Other Variations of Examples 6.1. To 6.6. Are Described Below

In one embodiment, the user uses his/her hands to press on the skin ofthe patient while inserting the cannula and injecting the material inorder to maintain the patient's tissues from the outside and to definethe localization of the material.

In one embodiment, the user injects the material using the same devicedescribed previously but injects the material in a bolus shape, which isexpanding the surrounding tissues of the injection site.

In one embodiment, the soft tissue engineering material is combined withthe graft of adipose tissues preliminary harvested from the patient. Forexample, adipose tissues are extracted using a harvesting cannula byliposuction. Tissues are let sediment for 5 minutes allowing to removethe blood and oil floating above the adipose tissues. The adiposetissues are distributed in 10 mL syringes. In one embodiment, the userinjects one spaghetti of adipose tissues of 2 mL to 10 mL using theColeman method and then he/she injects a spaghetti of the soft tissueengineering material. In another embodiment, the adipose tissues aremixed with the soft tissue engineering material by connecting twosyringes (one containing the adipose tissues, the other one containingthe soft tissue engineering material) using a Luer-to-Luer connector andby pushing sequentially on the two pistons of the two syringes. Themixture obtained is then injected using the method described before.

In another embodiment, the soft tissue engineering material is manuallydistributed in a body cavity (such as a breast cavity after siliconeimplant removal) using a sterile spatula in order to create a layer ofthe soft tissue engineering material.

In another embodiment, the soft tissue engineering material is suturedto surrounding tissues (in the case of large particles).

F) Clinical Results Obtained and Comparative Studies with Prior ArtMaterials

Example 7

A comparison of the stability and migration of 4 different materials,including the soft tissue engineering material according to theinvention was performed. The materials were the following:

“HA 1” is a commercially available hyaluronic acid based filler(“Juvederm Ultra 2” from Allergan.

“HA 2” is a commercially available, strongly crosslinked hyaluronic acidbased filler (“Macrolane” from Q-med AB).

“Matrix” is a commercially available, collagen based, flowable matrixused for wound repair (from Integra LifeSciences corporation).

“Material developed” is the material obtained in examples 1 to 4.

A defined volume of the different tested items (200 microliters) wasinjected subcutaneously in CD1 female mice in the back area of theanimal. Two samples of each tested item were injected, namely one oneach side of the spinal cord of the animal. In the case of the siliconeitem, the samples were implanted by first performing an incision in theskin of the animal and by inserting manually with tweezers the layer ofsilicone. The volumes of the items were monitored over time usingexternal measurements with a Caliper and using MRI scanning and MRIimages analysis. After 3 and 6 months, the animals were euthanized andhistology of the different implanted materials was performed.Bio-integration (percentage of the material occupied by cells andtissues, vascularization) was quantified. The results are represented inFIG. 9 , which shows macroscopic pictures of the implants site in micefor the different tested items, at different time points (t=0 is justafter the injection step). In the drawings on the left side of FIG. 9 ,the dashed lines represent the implant localization just after theinjection. The grey surface represents the implant localization 3 monthsafter the injection.

The macroscopic observation of the histology samples (see FIG. 10 )enabled to show that the material developed lead to the growth ofvessels and tissues, whereas the HA-based materials tested remainedclear and transparent, showing neither tissular nor vascular ingrowth.The

Matrix 1 material was populated with cells but it degraded before the 6months timepoint. These results are represented graphically in FIG. 11 .

The presence of a capsule surrounding the implants compared wasinvestigated on histological sections stained with Masson trichrome. Anadditional material was included in this comparative study, namely“Silicone” which a silicone layer sample cut from a silicone tissueexpander used in breast (Natrelle 133) Tissue expander from Allergan.The implanted samples were squares of 6 mm side and measured 1.5 mm inthickness.

The thickness of the capsule was measured for each material tested. Theresults are presented in Table 4 below:

TABLE 4 Material Material Matrix implanted developed HA1 HA2 1 SiliconeThickness of No No 92 +/− 17 No 104 +/− 20 the capsule capsule capsulemicrometers capsule micrometers (3 months after the implantation

It was observed that the soft tissue engineering material according tothe invention was stable over time. It did not migrate or increase involume. On the contrary, HA 1 was increasing in volume and the twoimplants merged together (they moved from the initial position). HA 2was also stable but the histological analysis showed the presence of aforeign body reaction (thin capsule around the implant, presence ofgiant cells), which could explain the stability of the position. Thematerial was isolated from the body and did not migrate. The matrix 1material was rapidly resorbed and did not produce a durable volume.

Although the invention has been described in conjunction with specificembodiments thereof, it is evident that many alternatives, modificationsand variations will be apparent to those skilled in the art.Accordingly, it is intended to embrace all such alternatives,modifications and variations that fall within the scope of the appendedclaims.

It is appreciated that certain features of the invention, which are, forclarity, described in the context of separate embodiments, may also beprovided in combination in a single embodiment. Conversely, variousfeatures of the invention, which are, for brevity, described in thecontext of a single embodiment, may also be provided separately or inany suitable subcombination or as suitable in any other describedembodiment of the invention. Certain features described in the contextof various embodiments are not to be considered essential features ofthose embodiments, unless the embodiment is inoperative without thoseelements.

1-55. (canceled)
 56. Implantable soft tissue engineering materialcomprising a multitude of particles wherein the implantable soft tissueengineering material is in form of a malleable paste and has adeployment pressure of above about 30 Pa, wherein each of the particlesis formed from a three-dimensionally warped and branched sheet that hasa mean sheet thickness (T), wherein the sheet comprises a biocompatiblematerial, wherein the particles have irregular shapes and comprise anumber of protrusions, wherein the particles have a number ofinterconnected channel-type conduits: wherein the conduits have a meandiameter (DC); and wherein the ratio of DC/T is larger than
 1. 57. Theimplantable soft tissue engineering material according to claim 56,further comprising one or more of the following substances:physiologically acceptable fluid, physiological saline, phosphatebuffered saline, blood plasma, living aspirates, lipoaspirate, water,aqueous solution, blood, serum, pharmaceutically active agents,lidocaine, adrenaline, cell suspensions, biological tissues, stem cells,virus, bacteria, fungi, transfecting agents, antibodies, geneticallymodified cells, extracellular matrices, co-cultures of cells, growthfactors, platelet rich plasma, cell differentiation factors, lipids andhigh density lipoprotein (HDL).
 58. The implantable soft tissueengineering material according to claim 56, characterized in that themultitude of particles is admixed with adipose tissue, engineeredtissue, autologous tissues, heterologous tissues, and/or xenologoustissues.
 59. The implantable soft tissue engineering material accordingto claim 56, characterized in that it is reversibly compressible afterinjection into a patient by uptaking liquid from the surroundingtissues.
 60. (canceled)
 61. The implantable soft tissue engineeringmaterial according to claim 56, characterized in that it exhibits aYoung's modulus between 100 Pa and 15 kPa, preferably between 200 Pa and5 kPa. 62-75. (canceled)
 76. The implantable soft tissue engineeringmaterial according to claim 56, wherein the implantable soft tissueengineering material has a deployment pressure of at least about 100 Pa.77. The implantable soft tissue engineering material according to claim56, wherein the biocompatible material of the sheet is selected from thegroup consisting of poly-ethyleneglycol (PEG), poly-acrylamide,poly-(hydroxyethyl)methacrylate, gelatin, and polysaccharides.
 78. Theimplantable soft tissue engineering material according to claim 56,wherein the biocompatible material of the sheet is a polysaccharideselected from the group consisting of hyaluronic acid, alginate,agarose, chitosan, cellulose, methylcellulose, carboxymethylcellulose,agarose, polysucrose, and dextran.
 79. The implantable soft tissueengineering material according to claim 56, wherein the biocompatiblematerial of the sheet is a material selected from the group consistingof carbohydrates, collagens, peptides, and extracellular matrices. 80.The implantable soft tissue engineering material according to claim 56,wherein the biocompatible material of the sheet is a hydrogel.
 81. Theimplantable soft tissue engineering material according to claim 56,wherein the implantable soft tissue engineering material deploys up to apredefinable and controllable volume via the control of the hydrationlevel of the particles.
 82. The implantable soft tissue engineeringmaterial of claim 56, wherein the particles are hydrated and preferablycomprise at least 0.05 weight % of the biocompatible material based onthe total weight of the hydrated particles, preferentially between 0.05weight % and 15 weight %, preferentially between 0.1 weight % and 5weight %.
 83. The implantable soft tissue engineering material of claim56, wherein the particles comprise a composite of collagen and apolysaccharide, optionally wherein the collagen is covalently attachedto the polysaccharide.
 84. A shapeable tissue or organ body implantcomprising the implantable soft tissue engineering material of claim 56.85. The shapeable tissue or organ body implant of claim 84 in a formsuitable for breast tissue reconstruction.
 86. The shapeable tissue ororgan body implant of claim 84 in a form suitable for treating tissuedefects.
 87. The shapeable tissue or organ body implant of claim 84 in aform suitable for lipofilling.
 88. The shapeable tissue or organ bodyimplant of claim 84 in a form suitable for aesthetic restoration in apatient's face or part of a patient's body, optionally wherein the partof the body is the breast.
 89. A method of implanting the implantablesoft tissue engineering material of claim 56 in a patient, comprising(i) placing the material in contact with a biological material to form amixture, and (ii) subsequently injecting the mixture in a target area ina patient.
 90. The method of claim 89, wherein the biological materialis adipose tissue, and wherein the method further comprises prior tostep (iii), (iv) mixing the material with the adipose tissue using twosyringes and a connector.
 91. The method of claim 90, wherein theadipose tissue is harvested from the patient.
 92. The method of claim89, wherein following step (iii), (v) the implant is manually shapedfrom outside the patient to create the desired shape.